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Gabriela Ferreira de Vasconcelos Martins
Mestre em Química
New devices to monitor oxidative stress
biomarkers in point-of-care: a new tool
for cancer prevention
Dissertação para obtenção do Grau de Doutor em
Nanotecnologias e Nanociências
Orientador: Maria Goreti Sales, Professor Adjunto, Instituto Superior de Engenharia do Porto, Instituto Politécnico do Porto Co-orientador: Elvira Maria Correia Fortunato, Professor Catedrático, Faculdade de Ciências e Tecnologia, Universidade NOVA de Lisboa
Júri:
Presidente: Prof. Doutor Rodrigo Ferrão de Paiva Martins Arguentes: Prof. Doutor Arben Merkoçi Prof. Doutor Carlos Manuel de Melo Pereira Vogais: Prof. Doutor Rodrigo Ferrão de Paiva Martins Prof. Doutor Francisco Miguel Gama Prof. Doutor Hugo Manuel Brito Águas Prof. Doutora Maria Goreti Ferreira Sales Prof. Doutor Bruno Costa-Silva
Dezembro de 2018
Gabriela Ferreira de Vasconcelos Martins
Mestre em Química
New devices to monitor oxidative stress
biomarkers in point-of-care: a new tool
for cancer prevention
Dissertação para obtenção do Grau de Doutor em
Nanotecnologias e Nanociências
Orientador: Maria Goreti Sales, Professor Adjunto,
Instituto Superior de Engenharia do Porto, Instituto
Politécnico do Porto
Co-orientador: Elvira Maria Correia Fortunato, Professor
Catedrático, Faculdade de Ciências e Tecnologia,
Universidade NOVA de Lisboa
Dezembro de 2018
ii
iii
New devices to monitor oxidative stress biomarkers in point-of-care: a new tool for
cancer prevention
Copyright: Gabriela Ferreira de Vasconcelos Martins
FCT/UNL e UNL
A Faculdade de Ciências e Tecnologia e a Universidade Nova de Lisboa têm o direito, perpétuo
e sem limites geográficos, de arquivar e publicar esta dissertação através de exemplares
impressos reproduzidos em papel ou de forma digital, ou por qualquer outro meio conhecido ou
que venha a ser inventado, e de a divulgar através de repositórios científicos e de admitir a sua
cópia e distribuição com objectivos educacionais ou de investigação, não comerciais, desde
que seja dado crédito ao autor e editor.
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Acknowledgments
Like someone great said one day "Above all, do not fear difficult moments, the best comes from
them..." (Dra. Rita Levi-Montalcini). This project and thesis is the outcome not only of my
individual work, but also it is the reflection of all the important people that surrounded me and to
whom I would like to thank.
Firstly, I want to acknowledge my supervisors, Prof. Goreti Sales and Prof. Elvira Fortunato for
giving me the opportunity to develop this work, under their close guidance. They are both role
models in science, and life, that had inspired me each day to never give up, searching with
imagination a solution to every problem. To Prof. Goreti I deeply thank the daily
encouragement, the trust in my choices, the scientific advices in our discussions and the
constant enthusiasm that she always bring to our "science talks". I thank to Prof. Elvira sharing
her scientific knowledge, rising important questions in our meetings, that in most cases ended
with great achievements in our work. Also, I want to acknowledge Prof. Elvira the opportunity to
develop scientific work in the CENIMAT facilities, whereas it was possible to explore and
investigate extra scientific approaches. Thank you to both for the support, confidence and
supervision along this work.
To Fundação para a Ciência e Tecnologia (FCT), the financial support that enabled the
development of this scientific work.
To all my colleagues in the BioMark that, in some way, contributed to this work; specially, Ana
Patrícia for our enthusiastic "electro-talks", Helena Gomes for her advices and good sense,
Liliana Truta, Felismina Moreira, Alexandra Santos, Carolina Hora and Mariana Carneiro for
their patience, support and kindness and all others colleagues.
To Ana Marques, from FCT-UNL, the help in the lab and the hospitality during my stay in
Lisbon, as well as the scientific comments along my thesis work.
To Dr. Rui Fernandes (TEM analysis, I3S) and Dr. Rui Rocha (SEM analysis, CEMUP), all the
constructive discussions related with the analysis of my samples.
To all my friends that have been my support, being always present and available when I need it
(because science is full of ups and downs). To Esther Garcia, Ana Torres and Sofia Caridade
thank you for continuing this crazy wonderful friendship that, despite the long distance, remains
always with me.
To my family, for their presence in my life, their support in my decisions and their kindness in
the right moments... Specially, to my parents for always trusting in me, knowing that I am a
reflection of them; to my brother, for being there when I need it (even without knowing it); to my
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grandmother Emília for being an inspiration in every day of my life, she is the true role model in
my beliefs and she will always continue to be...
To my dear husband Ricarte, for truly believing in me when I couldn't, for being my constant
support and most enthusiastic admirer. I wouldn't be here without you!
And finally, to my best achievement in life, to my dear incredible son Guilherme, that since the
first day of life inspires me to be each day better and stronger... I hope that science will also
inspire him one day!
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Resumo
Segundo as recentes estatísticas da Organização Mundial de Saúde (OMS), a segunda
principal causa de morte a nível mundial é o cancro, com uma taxa de mortes em 2018 de 9.6
milhões de pessoas. Em particular, as doenças relacionadas com cancro já originaram 26% do
número total de mortes ocorridas em Portugal em 2016. Vários mecanismos estão associados
ao desenvolvimento de cancro, sendo que o stress oxidativo parece exercer um papel crucial
na origem desta doença. Assim, a deteção precoce de múltiplos biomarcadores do stress
oxidativo constitui uma ferramenta essencial na prevenção do cancro e também na seleção das
terapias mais eficazes.
A procura de metodologias para análise específica de biomarcadores do stress oxidativo in loco
continua a ser um desafio para a investigação biomédica. Até ao momento, os métodos
analíticos utilizados para o diagnóstico de cancro, que incluem um exame patológico, são
insuficientes para deteção precoce da progressão do tumor. Assim, para ultrapassar esta
necessidade, o principal objetivo deste projeto é desenvolver métodos de deteção rápidos,
simples e precisos para quantificação de biomarcadores de stress oxidativo, com metodologias
de recolha não-invasivas, de modo a conduzir a um diagnóstico rápido e de confiança numa
fase inicial da doença.
Para este efeito, esta tese apresenta o fabrico de biomateriais com propriedades sensoriais
integradas em substratos condutores inovadores, para deteção in loco de biomarcadores de
stress oxidativo. De modo a obter processos de reconhecimento bioquímico de elevada
seletividade e especificidade, foi utilizada uma tecnologia de impressão molecular, que permite
criar locais artificiais de reconhecimento. No decorrer da fabricação das plataformas
transdutoras eletroquímicas, o papel foi usado como material de suporte alternativo aos
materiais convencionais geralmente incorporados nos sistemas de elétrodos.
Em suma, espera-se que os resultados deste plano possam contribuir, no futuro, para o
desenvolvimento e aplicação de plataformas de multi-analitos para rápida e simultânea deteção
de biomarcadores do stress oxidativo num contexto local.
Termos-chave: Biosensor eletroquímico; Polímero de impressão molecular; Biomarcador do
stress oxidativo; Deteção no local; Substrato de papel.
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Abstract
According to the most recent World Health Organization (WHO) data, cancer is the second
leading cause of death worldwide, accounting for 9.6 million deaths in 2018. In particular,
cancer diseases have caused 26% of the total deaths in Portugal in 2016. Among the complex
mechanisms associated to cancer development, Oxidative Stress (OS) seems to play an
important role at the origin of the disease. Thus, early diagnosis of multiple OS biomarkers may
be a fundamental tool in cancer prevention and in more efficient therapeutic strategies.
Despite the development and the research efforts that are being made, accurate and early
detection methods for cancer are still lacking. The demand for specific OS biomarker assays
carried out in wide screening programs in point-of-care (POC) is undoubtedly a difficult but
potentially useful challenge for biomedical research and health. So far, current methods for
cancer diagnosis based upon pathological examination alone are insufficient for detecting early
tumour progression.
Thus, to overcome this need, the present project aims the development of quick, simple and
accurate detection of selected OS biomarkers, collected using minimally invasive methods, in
order to allow rapid and reliable diagnosis at early stages of the disease. Under this scope, the
design of sensitive biosensing materials integrated with novel conductive substrates for POC
screening of OS biomarkers will be presented. In order to achieve a specific and highly selective
bio-chemical recognition process, molecular imprinting strategy was used to create the artificial
recognition sites. During the fabrication of electrochemical transduction platforms, paper was
introduced as a novel alternative to the conventional support materials usually incorporated in
electrode systems.
Overall, it is expected that the outcome of this plan will contribute, in the future, to the
development and application of a multi-analyte platform for simultaneous fast screening of
cancer biomarkers in POC context.
Keywords: Electrochemical biosensor; Molecular imprinting polymer; Oxidative stress
biomarker; Point-of-care sensing; Paper substrate.
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Table of Contents
1 FRAMEWORK ....................................................................................................................... 1
1.1 MOTIVATION .................................................................................................................. 1
1.2 OBJECTIVES .................................................................................................................. 2
1.3 THESIS OUTLINE ........................................................................................................... 2
1.4 LIST OF PUBLICATIONS ............................................................................................... 4
1.4.1 Papers published in international scientific journals ................................................ 4
1.4.2 Communications presented in national and international scientific conferences .... 4
2 INTRODUCTION .................................................................................................................... 7
2.1 OXIDATIVE STRESS ........................................................................................................... 7
2.1.1 Biomarkers of Oxidative Stress ............................................................................. 10
2.1.1.1 8-Hydroxy-2'-deoxyguanosine ........................................................................... 12
2.1.1.2 3-Nitrotyrosine .................................................................................................... 13
2.1.1.3 Malondialdeyde and Hydroxynonenal ................................................................ 15
2.2 BIOSENSORS .............................................................................................................. 16
2.2.1 Recognition Element.............................................................................................. 17
2.2.2 Signal Transduction ............................................................................................... 17
2.3 ELECTROCHEMICAL BIOSENSORS ..................................................................................... 19
2.3.1 Voltammetry ........................................................................................................... 20
2.3.2 Cyclic Voltammetry ................................................................................................ 20
2.3.3 Square Wave Voltammetry .................................................................................... 22
2.3.4 Differential Pulse Voltammetry .............................................................................. 22
2.3.5 Electrochemical Impedance Spectroscopy............................................................ 23
2.3.6 Electrode size, materials and supports .................................................................. 26
2.4 MOLECULAR IMPRINTING POLYMER ....................................................................... 33
2.5 NANOMATERIALS ....................................................................................................... 38
3 8-HYDROXY-2'-DEOXYGUANOSINE BIOMARKER DETECTION DOWN TO
PICOMOLAR LEVEL ON A PLASTIC ANTIBODY FILM .......................................................... 43
3.1 INTRODUCTION .......................................................................................................... 44
3.2 EXPERIMENTAL SECTION ......................................................................................... 45
3.2.1 Reagents and Materials ......................................................................................... 45
3.2.2 Apparatus .............................................................................................................. 45
3.2.3 Gold electrode cleaning ......................................................................................... 45
3.2.4 Sensor fabrication .................................................................................................. 46
xii
3.2.5 Electrochemical assays ......................................................................................... 46
3.2.6 Surface analysis .................................................................................................... 47
3.2.7 Preparation and characterization of the FITC-labeled surfaces ............................ 47
3.2.8 Selectivity studies and analysis in urine samples .................................................. 47
3.3 RESULTS AND DISCUSSION ...................................................................................... 48
3.3.1 Optimization of experimental variables.................................................................. 48
3.3.2 Preparation and electrical follow-up of MIP sensor ............................................... 50
3.3.3 Characterization of the modified surfaces ............................................................. 53
3.3.4 Performance of MIP sensor ................................................................................... 55
3.3.4.1 Calibration curve ................................................................................................ 55
3.3.4.2 Selectivity studies ............................................................................................... 56
3.3.4.3 Analysis of spiked human urine samples ........................................................... 57
3.4 CONCLUSIONS ............................................................................................................ 58
4 PAPER-BASED SENSING DEVICE FOR ELECTROCHEMICAL DETECTION OF
OXIDATIVE STRESS BIOMARKER 8-HYDROXY-2'-DEOXYGUANOSINE IN POINT-OF-
CARE .......................................................................................................................................... 61
4.1 INTRODUCTION .......................................................................................................... 62
4.2 EXPERIMENTAL SECTION ......................................................................................... 63
4.2.1 Reagents and Materials ......................................................................................... 63
4.2.2 Apparatus .............................................................................................................. 63
4.2.3 Fabrication and characterization of the paper-based sensor ................................ 64
4.2.4 Electrochemical assays ......................................................................................... 64
4.3 RESULTS AND DISCUSSION ...................................................................................... 65
4.3.1 Electrochemical behaviour of 8-OHdG .................................................................. 65
4.3.2 DPV analysis of 8-OHdG on paper-modified electrodes ....................................... 66
4.3.3 Characterization of the paper-modified electrodes ................................................ 67
4.3.4 Optimization of DPV experimental conditions ....................................................... 68
4.3.5 Analytical applications ........................................................................................... 70
4.3.5.1 Calibration curve ................................................................................................ 70
4.3.5.2 Selectivity ........................................................................................................... 71
4.3.5.3 Serum samples .................................................................................................. 73
4.4 CONCLUSIONS ............................................................................................................ 75
5 NOVEL WAX-PRINTED PAPER-BASED DEVICE FOR A DIRECT ELECTROCHEMICAL
DETECTION OF 3-NITROTYROSINE ........................................................................................ 77
5.1 INTRODUCTION .......................................................................................................... 78
5.2 EXPERIMENTAL SECTION ......................................................................................... 79
5.2.1 Reagents and Materials ......................................................................................... 79
xiii
5.2.2 Fabrication of paper-based SPE ........................................................................... 80
5.2.3 Electrochemical assay ........................................................................................... 81
5.2.4 Surface characterization of the paper-based SPEs .............................................. 81
5.2.5 Detection of 3-Nitrotyrosine onto the paper-based SPE ....................................... 81
5.2.6 Selectivity assay .................................................................................................... 81
5.3 RESULTS AND DISCUSSION ...................................................................................... 82
5.3.1 Electrochemical performance of the paper-based SPEs....................................... 82
5.3.2 Morphological characterization of the paper-based SPEs .................................... 85
5.3.3 Direct detection of 3-Nitrotyrosine ......................................................................... 86
5.3.4 Calibration and interference assay ........................................................................ 87
5.4 CONCLUSIONS ............................................................................................................ 89
6 ELECTROCHEMICAL PAPER-BASED BIOSENSOR FOR LABEL-FREE DETECTION
OF 3-NITROTYROSINE IN HUMAN URINE SAMPLES USING MOLECULAR IMPRINTED
POLYMER ................................................................................................................................... 91
6.1 INTRODUCTION .......................................................................................................... 92
6.2 EXPERIMENTAL SECTION ......................................................................................... 94
6.2.1 Reagents and Materials ......................................................................................... 94
6.2.2 Apparatus .............................................................................................................. 94
6.2.3 Electrochemical assay ........................................................................................... 94
6.2.4 Assembly of the imprinted-based biosensor ......................................................... 95
6.2.5 Analysis of urine samples ...................................................................................... 95
6.3 DISCUSSION AND RESULTS ...................................................................................... 96
6.3.1 Electrochemical study ............................................................................................ 96
6.3.2 Electropolymerization of phenol - MIP versus NIP ................................................ 98
6.3.3 Optimization of experimental conditions during MIP assembly ........................... 100
6.3.3.1 Effect of scan-rate and number of electropolymerization cycles ..................... 100
6.3.3.2 Effect of monomer concentration ..................................................................... 101
6.3.3.3 Effect of imprinted 3-NT concentration ............................................................ 102
6.3.4 Characterization of the modified paper-electrodes .............................................. 104
6.3.5 Performance of the imprinted-sensor .................................................................. 105
6.3.5.1 Calibration curve .............................................................................................. 105
6.3.5.2 Urine samples .................................................................................................. 106
6.4 CONCLUSIONS .......................................................................................................... 107
7 SYNTHESIS AND CHARACTERIZATION OF CORE-SHELL MAGNETIC
NANOPARTICLES ................................................................................................................... 109
7.1 INTRODUCTION ........................................................................................................ 109
7.2 EXPERIMENTAL SECTION ....................................................................................... 111
xiv
7.2.1 Reagents and Materials ....................................................................................... 111
7.2.2 Apparatus ............................................................................................................ 112
7.2.3 Synthesis of core-shell nanoparticles .................................................................. 112
7.2.3.1 Fabrication of iron oxide nanoparticles ............................................................ 112
7.2.3.2 Preparation of iron oxide-silica core-shell ........................................................ 113
7.2.4 Characterization of the modified nanoparticles ................................................... 113
7.2.5 Electrochemical assays ....................................................................................... 114
7.3 DISCUSSION AND RESULTS .................................................................................... 114
7.3.1 Synthesis of iron oxide nanoparticles .................................................................. 114
7.3.2 Fabrication of iron oxide-silica core-shell ............................................................ 115
7.3.2.1 Optimization of the sol-gel process .................................................................. 119
7.3.2.2 Functionalization with APTES .......................................................................... 121
7.3.3 Characterization of the modified-nanoparticles ................................................... 123
7.4 CONCLUSIONS .......................................................................................................... 124
8 CONCLUSIONS AND FUTURE PERSPECTIVES ........................................................... 127
xv
List of Figures
Figure 2.1: Adapted scheme highlighting the various activators and inhibitors factors associated
to the production of ROS ............................................................. ..................................................8
Figure 2.2: Adapted scheme showing the proposed mechanisms for ROS production and their
contribution to the aging process......... .................................................................................... .....9
Figure 2.3: Proposed mechanism for the formation of 8-OHdG........... ..... ................................12
Figure 2.4: Adapted scheme showing the different pathways responsible for the nitration of
tyrosine residues of proteins and consequent production of 3-NT.................... .. .......................14
Figure 2.5: Chemical structure of A) malondialdehyde and B) 4-hydroxynonenal...... ........... ....15
Figure 2.6: Schematic representation illustrating the main constitution, function and practical
application of a biosensor device............................... ............................... ..................................16
Figure 2.7: Schematic representation with the main features of the different types of sensor
transduction............................................ ................... ..................................................................18
Figure 2.8: Cyclic voltammogram obtained for a reversible system......... ............. ....................21
Figure 2.9: Comparison of typical voltammograms obtained for reversible and irreversible
systems.............................................................................. .... .....................................................22
Figure 2.10: Nyquist plot for an electrochemical Faradaic system.................. .. ........................25
Figure 2.11: A typical example of a Bode plotting.......................... ............... .............................25
Figure 2.12: Main constituents of a paper-based assembly with three-integrated electrodes. . .27
Figure 2.13: Examples of the wide diversity of SPEs commercially available...... ........ .............27
Figure 2.14: Representation of the different field of applications concerning carbon-based
materials............................................................................... ..................... .................................28
Figure 2.15: Adapted scheme with the various routes available to pattern paper-based sensors,
with respective main advantages and limitations............... ....................... ..................................30
Figure 2.16: Schematic representation about techniques of nanofabrication, detection
methodologies and practical applications of paper sensors........... ............ ................................31
Figure 2.17: Some examples of different paper-based biosensors for A) virus, B) bacteria, C)
cell and D) multi-protein detection....................... ............... .........................................................32
Figure 2.18: Schematic representation of the synthesis of molecularly imprinted polymers.. . ..33
Figure 2.19: Graphical representation for publications under the issue A) "MIP" (molecularly-
imprinted polymer* or molecular imprinting or MIP*) and B) "biosensor+MIP" (biosensor* and
molecularly-imprinted polymer* or molecular imprinting or MIP*), from April 2018 ISI Web of
Knowledge................................................. ..................................................................................34
Figure 2.20: Proposed mechanism of phenol electro-oxidation............ .......... ...........................37
Figure 2.21: Representation of some nanostructured materials used for diagnostic
applications............................................... ...................... ............................................................38
Figure 2.22: Adapted scheme of the different states during nanoparticles fabrication...............39
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Figure 2.23: Different nanostructures of carbon A) graphene, B) SWCNTs and C)
MWCNTs.....................................................................................................................................39
Figure 2.24: Different immobilization methodologies used during the fabrication of a sensor
device: A) sandwich immunoassay approach; B) MIP-based approach; C) labeled nanoparticle
approach and D) ink-based approach.......................... .... ...........................................................40
Figure 2.25: Adapted graphic concerning the different routes used for the synthesis of iron
oxide magnetic nanoparticles.................. ................ ....................................................................41
Figure 2.26: Silica applications conjugated with magnetic nanoparticles as nanoplatforms... .. 42
Figure 3.1: Schematic representation of the assembly of the gold-modified imprinted
sensor............................... ..................... ......................................................................................44
Figure 3.2: A) Cyclic voltammograms of a gold-modified electrode immersed in 0.01 M PBS
aqueous solution containing different concentrations of monomer phenol (0.25, 0.5 and 1.25
mM), pH 7.4, scan rate 20 mVs-1
; B) Charge variation during electropolymerization of phenol (3
cycles) obtained from MIPs with different ratios of template to monomer (1:3 and 1:1) and NIP in
0.01 M PBS........ ........................................................................................................... ..............48
Figure 3.3: Calibration curves of 8-OHdG obtained for MIPs with different ratios of template/
monomer, 1:3 (closed gray circles) and 1:1 (open black circles)................. ........... ....................49
Figure 3.4: Cyclic voltammograms concerning the electropolymerization of 0.25 mM phenol in
0.01 M PBS, pH 7.4, (scan rate 20 mVs-1
, 3 cycles) at gold-modified electrodes with (dashed
line) and without (straight line) the template molecule 8-OHdG........ ........................... ..............50
Figure 3.5: A) CV of the gold electrode (green line), thiol-modified gold electrode (red line), NIP
and MIP after electropolymerization (blue and grey line, respectively) and after template
removal (black lines, on the right side), measured in aqueous solution containing 5 mM
[Fe(CN)6]3-/4-
in 0.01 M PBS pH 7.4 and B) EIS of (a) gold electrode, (b) thiol-modified gold
electrode, NIP (c) before and (d) after removal, MIP (e) before and (f) after removal, in aqueous
solution containing 5 mM [Fe(CN)6]3-/4-
in 0.01 M PBS................. ................ ..............................51
Figure 3.6: A) FTIR-ATR spectra of gold, thiol-modified gold, NIP and MIP electrodes; B)
RAMAN spectra of gold, thiol-modified gold, NIP and MIP electrodes and C) typical image from
RAMAN, measured at 50x magnification, of the gold-screen printed electrodes (Au-SPE).... . ..53
Figure 3.7: A) SEM micrographs of NIP and MIP electrodes and B) confocal imaging of FITC
antibody against 8-OHdG attached to NIP and MIP surfaces...................... ...............................54
Figure 3.8: A) Nyquist plot of MIP sensor in 5 mM [Fe(CN)6]3-/4-
in 0.01 M PBS pH 7.4,
previously incubated in increasing concentrations of 8-OHdG and B) the corresponding
calibration curves for both MIP and NIP sensors; C) Calibration curves of NIP and MIP sensors
for different 8-OHdG concentrations in urine samples, measured in 5 mM [Fe(CN)6]3-/4-
in 0.01
M PBS pH 7.4. All error bars represent the standard deviation for three independent
measurements............................................. ........... ....................................................................55
xvii
Figure 3.9: EIS measurement of MIP-based sensor recorded after incubation in 5 pg/mL 8-
OHdG solution, alone and in the presence of uric acid (0.4 g/mL), citric acid (0.5 g/mL) and
glucose (0.1 mg/mL). All solutions were prepared freshly on PBS pH 7.4.......... ........ ...............57
Figure 3.10: Calibration curve of 8-OHdG in a urine sample. Rct relative corresponds to the
normalized value of charge transfer resistance against the PBS measurement for each spiked
level and S is the slope of the experimental calibration, obtained from three independent
measurements........................ .................. ...................................................................................57
Figure 4.1: Schematic representation of the oxidation process of 8-OHdG molecule followed on
a conductive carbon paper substrate: 1) hydrophobic white paper as substrate; 2) conductive
carbon-coated paper; 3) in-situ electrochemical measurement........... ............. ..........................63
Figure 4.2: Successive cyclic voltammograms performed in PBS at pH 7.4 with 8-OHdG
molecule at different scan rates. Inset: calibration plot of the 8-OHdG oxidation peak current
versus scan rate............................ .......................... ....................................................................65
Figure 4.3: DPV detection of 200 ng/mL 8-OHdG solution in PBS pH 7.4 on different graphite-
based electrodes prepared after the incorporation of various nanomaterials dispersed in the
graphite ink, such as, PEDOT nanoparticles, CNTMW and CNTMW-COOH........ ......... ...........66
Figure 4.4: RAMAN spectra of the different graphite-based electrodes prepared after the
incorporation of nanomaterials dispersed in the graphite ink, such as, A1) PEDOT
nanoparticles, A2) Graphite, A3) CNTMW and A4) COOH-CNTMW, with the calculated ID/IG
ratios and B) RAMAN spectra with the magnification of the D (Disorder) band, in full-scale
mode............................................... .................... ........................................................................67
Figure 4.5: Successive differential pulse voltammograms of 0.1 mg/mL 8-OHdG in PBS pH 7.4
recorded (A) without any application of conditioning potential and (B) with a conditioning
potential of +0.20V applied before each measurement..................... ............ .............................68
Figure 4.6: Dependence of the sensor response on the (A) pre-accumulation potential and (B)
time of accumulation during 8-OHdG oxidation in PBS pH 7.4.......... ............... ..........................69
Figure 4.7: (A) Cleaning effect (after CV in PBS pH 7.4) on the 8-OHdG detection by DPV
signal and (B) sensor regeneration after voltammetric cycles performed in PBS pH 7.4... .. ......69
Figure 4.8: Differential pulse voltammograms recorded for 8-OHdG solutions prepared in
different buffer solutions, with different pH values.............. .................... ...................................70
Figure 4.9: A) Differential pulse voltammograms for different concentrations of 8-OHdG
prepared in PBS pH 7.4 and (B) calibration plot of the concentration of 8-OHdG........ ....... .......71
Figure 4.10: (A) DPV recordings for individual solutions with concentrations of 0.1 mM of 8-
OHdG, ascorbic acid and uric acid in PBS at pH of 7.4; (B) DPV recording of a mixture with all
of the 3 compounds, in the same concentrations.......................... .............. ...............................72
Figure 4.11: Differential pulse voltammograms for serum samples diluted 1:10 in different
buffers, such as, (A) Tris pH 9.1, (B) PBS pH 7.4 and (C) Acetate pH 5.1, doped with 1 ug/mL of
8-OHdG......................................................................... .................. ............................................73
Figure 4.12: Calibration curve of the concentration of 8-OHdG in diluted serum samples..... . ..74
xviii
Figure 5.1: Schematic illustration of the different steps related to the sensor device, namely, A)
the electrochemical apparatus for biological samples assessment; B) photo and morphological
characterization of the paper-based electrodes; and C) the assembly of the electrochemical
sensing platform........................ .................. ................................................................................79
Figure 5.2: Detailed scheme of the fabrication of the paper-based electrodes........... .... ..........80
Figure 5.3: Cyclic voltammograms for 5 mM [Fe(CN)6]4-/3-
redox couple in A) 0.1 M KCl solution
and B) PBS pH 7.4, at different scan-rates; Plot representation of both the anodic and cathodic
peak currents versus the square-root of the scan-rate for 5 mM [Fe(CN)6]4-/3-
redox couple in C)
0.1 M KCl solution and D) PBS pH 7.4................................ ................ .......................................82
Figure 5.4: A) Cyclic voltammograms for 5 mM [Fe(CN)6]4-/3-
redox couple in 0.1 M KCl, at
different scan-rates and B) plot representation of both the anodic and cathodic peak currents
versus the square-root of the scan-rate for 5 mM [Fe(CN)6]4-/3-
redox couple...... ........ ..............83
Figure 5.5: Effect of the different redox probes upon the electrochemical response. A) plots the
peak potential separation (ΔE) versus the scan-rate and B) the anodic and cathodic peak
current ratio (IpA/IpC) versus the scan-rate for [Fe(CN)6]4-/3-
and [Ru(NH3)6]3+
probes at 5 mM
concentration in 0.1 M KCl; Plots of current peak versus the probe concentration for C)
[Fe(CN)6]4-/3-
and D) [Ru(NH3)6]3+
, at a scan-rate of 50 mV/s, in 0.1 M KCl and PBS pH 7.4. .. ..84
Figure 5.6: SEM images of the A) and B) WE carbon-surface at different magnifications and C)
and D) Cross-section imaging of the carbon-layer at different magnifications............ ...... .........85
Figure 5.7: CV recordings over the potential range -1 V to +1 V in 0.1 M phosphate buffer with
(colour line) and without (dashed line) 1 mM of 3-NT, at a scan-rate of 50 mV/s, and in the inset
figure the chemical structure of 3-NT.................................. .............. ..........................................86
Figure 5.8: SWV response of 1 mg/mL 3-NT A) in different supporting electrolyte solutions and
B1) in 0.1 M phosphate buffer solution at different pH values ranging from 6 to 8. B2) Plot of the
potential value of the SWV versus the pH obtained in 1 mg/mL 3-NT in phosphate buffer
solution...................................................................... ................... ...............................................87
Figure 5.9: A) SWV recordings of 3-NT at different concentrations, in 0.1 M phosphate buffer at
7.4 pH (inset figure is for lower concentrations) and B) calibration curve of 3-NT, with and
without the application of an accumulation potential................ ............. ......................................88
Figure 5.10: A) Electrochemical response of tyrosine, ascorbic acid, uric acid and creatinine
over the studied potential range and B) the curves of calibration for 3-NT only and in the
presence of 10 μM of tyrosine......................................... ........................ ....................................88
Figure 6.1: Illustration of the sensor film fabrication by molecular imprinting for recognition of 3-
nitrotyrosine.................................... ................. ............................................................................93
Figure 6.2: Cyclic voltammograms of 3-nitrotyrosine in A) PBS solution, at different scan
directions, over the potential range -1 V to +1 V; B) three different electrolyte solutions, over the
potential range -0.4 V to +1.2 V; and C) KCl solution, over the potential range +0.2 V to +0.8 V.
Cyclic voltammograms of D) phenol and 3-nitrotyrosine, individually and E) mixture phenol + 3-
nitrotyrosine. Chemical representation of F) phenol and G) 3-nitrotyrosine............. ........ ..........96
xix
Figure 6.3: Cyclic voltammograms of NIP and MIP electrodes during electrochemical
polymerization of phenol for the A) 1st and B) 5
th scan cycle, in KCl solution (0.1 M, pH 5.9). EIS
obtained for each step of the construction for C) NIP and D) MIP electrodes, in 5 mM solution of
K3[Fe(CN)6] and K4[Fe(CN)6] prepared in phosphate buffer solution (0.1 M, pH 6.0)............. ....98
Figure 6.4: A) EIS measurements obtained before and after template removal, at different
scan-rates: A) 15 mV/s; B) 50 mV/s and C) 150 mV/s and, with different number of cycles: D) 2;
E) 5 and F) 10, recorded during phenol electropolymerization..................... .............. ..............101
Figure 6.5: EIS obtained for NIP and MIP sensors at two different phenol concentrations, A) 1
mM and B) 0.25 mM. C) 3-Nitrotyrosine response for both MIP electrodes, obtained from DPV
measurements................................................ ......................... .................................................102
Figure 6.6: A) Charge variation during phenol electropolymerization (5 cycles) obtained from
MIPs with different concentrations of template molecule; B) EIS obtained for the MIPS with
different concentrations of template molecule; DPV measurements after contact with different
concentrations of 3-NT for MIPs with C) 0.05 mM, D) 0.25 mM and E) 0.50 mM concentration of
template molecule; F) Scheme related to the distribution of imprinting sites..... ........ ...............103
Figure 6.7: A) Raman spectra of clean carbon-based electrode, NIP and MIP-modified
surfaces. SEM images of B) NIP and C) MIP materials................. ........... ................................104
Figure 6.8: Calibration curves corresponding to the response of A) MIP and B) NIP sensors
against the concentration of 3-nitrotyrosine. The inset figure is related to the DPV recordings for
each standard concentration.................... .................. ...............................................................105
Figure 6.9: Calibration curves corresponding to the response of A) MIP and B) NIP sensors
against the concentration of 3-nitrotyrosine, performed in 1:10 diluted human urine
samples............................................................. ...................... ..................................................106
Figure 7.1: Schematic representation of the different steps related to the synthesis of the core-
shell magnetic nanoparticles............................................ ...................... ...................................110
Figure 7.2: Synthesis reaction of the iron-oxide nanoparticles via co-precipitation
method............................................................................... .............. .........................................113
Figure 7.3: (A-B) TEM images of iron-oxide nanoparticles at different magnifications and C)
image of the MNPs under the application of a magnetic field.......... ................. ........................115
Figure 7.4: Square-wave voltammograms concerning the two (individual) redox probes
ruthenium and NADH, in PBS at a pH 7.4, applied in a clean, bare carbon-SPE....... .... .........115
Figure 7.5: SEM images of the MNPs A) non-modified, modified with B) SiO2 with NADH and
C) SiO2 with ruthenium; EDS spectra of the MNPs D) non-modified, modified with E) SiO2 with
NADH and F) SiO2 with ruthenium......................................... ................. ..................................117
Figure 7.6: A) Cyclic voltammogram applied in PBS at pH 7.4 of the different types of MNPs; B)
images of the wet suspension of the MNPs B1) non-modified, modified with B2) SiO2 with NADH
and B3) SiO2 with ruthenium and TEM images of the MNPs C1) non-modified, modified with C2)
SiO2 with NADH and C3) SiO2 with ruthenium............................ ............... ...............................118
xx
Figure 7.7: TEM (A) and SEM (B) imaging of the silica-based nanoparticles prepared (1)
without ethanol and SDS, (2) without ethanol and with SDS and (3) with ethanol and
SDS.................................................................. ................... ......................................................119
Figure 7.8: Square wave voltammograms of the silica-based MNPs synthesized with increasing
concentration of TEOS A) 0.1 mL, B) 0.5 mL and C) 1.0 mL; TEM images of the silica-based
MNPs obtained with different TEOS concentration D) 0.1 mL, E) 0.5 mL and F) 1.0 mL ......... 120
Figure 7.9: Schematic illustration of the fabrication procedure of the core-shell magnetic
nanoparticles and their application as electrochemical probes ................................................. 121
Figure 7.10: Square wave voltammograms of the developed silica-based MNPs at different
stages of fabrication: A) TEOS with ruthenium, B) effect of a magnetic field on the previous
MNPs and C) the effect of functionalization with APTES; D) Image of the electrochemical
measurement of MNPs using a magneto .................................................................................. 122
Figure 7.11: FTIR spectra of the magnetic nanoparticles at each step of the fabrication and
modification reaction ................................................................................................................. 123
Figure 7.12: TG analysis of the different stages of the magnetic nanoparticles ...................... 124
xxi
List of Tables
Table 2.1: Summary of the most relevant OS biomarkers and their main categories ................ 11
Table 2.2: Comparison of paper as a substrate material with other traditional materials .......... 31
Table 3.1: Comparison of the main characteristics of some reported assays used in the
detection of 8-OHdG................................................. ..... .............................................................59
Table 4.1: Comparison of different electrochemical sensors for determination of 8-OHdG.. . ....76
Table 6.1: Comparison of the different sensors for 3-nitrotyrosine detection in biological
matrices................................................................. ............. .........................................................93
Table 6.2: Analytical data obtained from the Raman spectra related to the different
modifications........................................................... ......... .........................................................106
xxii
xxiii
List of Abbreviations
AFM Atomic force microscopy
APTES (3-Aminopropyl)triethoxysilane
ATR Attenuated total reflectance
C Capacitor
CE Counter/auxiliary electrode
CNTs Carbon nanotubes
CV Cyclic voltammetry
DMF Dimethylformamide
DNA Deoxyribonucleic acid
DPV Differential pulse voltammetry
EDS Energy dispersive X-ray spectroscopy
EIS Electrochemical impedance spectroscopy
ELISA Enzyme-linked immunosorbent assay
Fe2O3 Magnemite
α-Fe2O3 Hematite
-Fe2O3 Maghemite
FET Field-effect transistor
FITC Fluorescein isothiocyanate
FRP Free radical polymerization
FTIR Fourier-transform infrared spectroscopy
GC-MS Gas chromatography mass spectrometry
4-HNE 4-Hydroxynonenal
H2O2 Hydrogen peroxide
HPLC High performance liquid chromatography
ISFET Ion-selective field-effect transistor
L Inductor
LC-MS Liquid chromatography–mass spectrometry
LOD Limit of detection
MB Methylene blue
MDA Malonaldehyde
MIP Molecularly imprinted polymer
MNP Magnetic nanoparticle
MRI Magnetic resonance imaging
MWNT Multi-walled nanotube
MWNT-COOH Multi-walled nanotube, carboxylic acid functionalized
NIP Non-molecularly imprinted polymer
xxiv
NMR Nuclear magnetic resonance
NO Nitric oxide
NO2 Nitronium group
3-NT 3-Nitrotyrosine
O2 Oxygen
O2- Superoxide anion
O2-.
Superoxide radical
OH. Hydroxyl radical
8-OHdG 8-Hydroxy-2’-deoxyguanosine
8-OHG 8-Hydroxyguanosine
ONOO- Peroxynitrite
OS Oxidative stress
PBS Phosphate buffered saline
PEDOT Poly(3,4-ethylenedioxythiophene)
P3MT Poly (3-methylthiophene)
POC Point-of-care
PSA Prostate specific antigen
PVC Polyvinyl chloride
PVC-COOH Polyvinyl chloride, carboxylated
QCM Quartz crystal microbalance
R Resistor
RE Reference electrode
RNS Reactive nitrogen species
ROS Reactive oxygen species
RRS Resonance Raman spectroscopy
RSD Relative standard deviation
SDS Sodium dodecyl sulphate
SEM Scanning electron microscopy
SERS Surface-enhanced Raman spectroscopy
SiO2 Silicon dioxide
SnO2 Tin(IV) oxide
SPE Screen-printed electrode
SPR Surface plasmon resonance
SWNT Single-walled nanotube
SWV Square wave voltammetry
TBARS Thiobarbituric acid reactive substances
TGA Thermogravimetric analysis
TEM Transmission electron microscopy
TEOS Tetraethyl orthosilicate
UV-Vis Ultraviolet–visible spectroscopy
xxv
WE Working electrode
WHO World Health Organization
Symbols
Cd Double-layer capacitance
E Applied potential
E0 Amplitude of the voltage signal at t = 0
Epa Anodic peak potential
Epc Cathodic peak potential
ΔEp Difference between the anodic and cathodic peak potentials
I Current
Ipa Anodic peak current
Ipc Cathodic peak current
n Number of electrons involved in the electrochemical reaction
t Time
Rct Charge-transfer resistance
RΩ Solution resistance
V Potential
Z Impedance of the system
ZW Warburg impedance
θ Phase shift
ʋ Scan-rate
ω Angular frequency
xxvi
Chapter 1
1
CHAPTER 1
1 Framework
1.1 MOTIVATION
Currently, about 26% of Portuguese people died of cancer every year, according to the
most recent World Health Organization (WHO) data (WHO, 2016). As a leading cause of death
worldwide, the most common types of cancer death includes lung, liver, colorectal, stomach and
breast. Although there are some well-identified risk factors, such as, tobacco smoking, alcohol
consumption, obesity, among others, aging is another important factor to take in account for the
development of cancer. So, as the way a person grows older, the incidence of cancer rises
dramatically due to a danger combination of the risks for specific cancers with a less efficient
cellular repair mechanism [1][2].
In this context, cancer mortality can be substantially reduced if cases are detected and treated
early. However, despite all the research efforts that have been made in the last decades,
screening programmes for quick, low-cost and in loco diagnostic are still lacking. So, the
application of biochemical markers for the diagnosis and management of patient's status has
been a growing approach in recent years, while developments in molecular biology lead to a
continuous discovery of new circulating biomarkers. In parallel, molecular epidemiological
studies have evidenced a link between oxidative stress (OS) and carcinogenesis. Indeed, OS-
based biomarkers have been proven essential in revealing how OS may mediate toxicity
induced by many known carcinogenic environmental agents [3]. Looking closer to the sub-
products originated by OS, the most used and promising biomarkers found until now include 8-
hydroxy-2’-deoxyguanosine (8-OHdG), 3-nitrotyrosine (3-NT), malonaldehyde (MDA) and 4-
hydroxynonenal (4-HNE).
Nevertheless, the current analytical strategies used for the detection, monitoring and diagnosis
of cancer are still insufficient. Despite their high sensitivity and selective response, these hold
some limitations, such as, long analysis time, complex operation and large volume samples.
Alternatively, biosensors are miniaturized devices, portable, easy-to-use with a straightforward
operation, that can track in real-time multiple biomarkers in clinically relevant samples. Under
this purpose, a modified-transducer surface is fabricated herein and afterwards, functionalized
and tailored with suitable bioreceptor elements. So, a biomimetic polymeric network can be
Chapter 1
2
finely designed through specific interactions between the building blocks of a biocompatible
matrix and the desired specific target [4], enabling in the end, a specific, simple, inexpensive
biosensor device. Moreover, to accomplish a suitable performance regarding the selectivity and
sensitivity parameters of the sensor device, suitable electrochemical platforms that include
cellulose paper are designed and functionalized in order to fabricate facile point-of-care (POC)
testing device.
1.2 OBJECTIVES
The study and development of biosensor devices and novel supporting materials have
been the core-pieces of BioMark and CENIMAT research groups receiving this plan, including
the detection of cancer and neurodegenerative diseases. Under this context, the main goals of
this work are:
(1) Identification and selection of the OS biomarkers that are most relevant in cancer diseases;
(2) Design and integration of a bioreceptor material, especially using molecular imprinting
technology, directly assembled on an electrochemical sensing platform;
(3) Fabrication of low cost paper-based innovative conductive substrates to be fully-integrated
as biosensing platforms;
(4) And ultimately, the fabrication of nanostructured electrochemical labels to be incorporated, in
the future, in the development of a suitable multi-analyte platform for a simultaneous screening
in POC of relevant OS biomarkers;
1.3 THESIS OUTLINE
This thesis is organized in eight chapters.
Firstly, Chapter 1 describes the motivation and the main objectives of this work, with a
summary description of the structure of the thesis. The list of publications and communications
(oral and poster) associated to this PhD thesis is also presented.
Chapter 2 provides some fundamental background regarding the design and principle of
biosensor devices, emphasizing those with special relevance to the present project. A brief
state-of-the-art applied to the selected biomarkers is also described. An overview of the most
common used nanomaterials and their main applications is addressed herein.
Chapter 3 presents the main results concerning the development, characterization and
application of a biomimetic biosensor for electrochemical detection of 8-OHdG biomarker. This
oxidized base is the most widely used biomarker for sensing oxidative DNA damage. The
assembly of a plastic antibody film on the surface of a commercial gold electrode enabled the
detection down to picoMolar level. Herein, a conventional three-electrode system was chosen to
study the electrochemical performance of the plastic antibody designed towards 8-OHdG
Chapter 1
3
quantification. This configuration was suitable is terms of analytical features and sample
analysis feasibility, but it was not compatible with POC analysis.
In order to overcome the limitations of the conventional commercially-available electrodes,
Chapter 4, describes the design, fabrication and application of an innovative paper-based
sensing device for assessing 8-OHdG. Under this scope, paper was employed as an
environmental-friendly alternative to other electrode supports. Although the levels of detection
were not comparable with the previous commercial approach, the proposed carbon-based
electrochemical sensor hold special features, such as, the potential to be miniaturized into
smaller portable size, being disposable and low-cost. Still, this configuration required however
additional improvements in terms of production, as it was being produced solely by manual
techniques, which hindered the reproducibility of the different electrodes.
Thus, more advanced techniques of microfabrication were introduced and, Chapter 5 focused
on the production, characterization and application of a wax-printed paper-based device
enabling the integration of the 3 electrode in the same platform for the direct assessment of 3-
NT biomarker. Besides the novelty of using a flexible paper-based printed electrode, herein 3-
NT was chosen as the target molecule due to the known correlation between this product of
protein oxidation and many acute and chronic diseases. Although it is the first paper-based
electrochemical sensor to the detection of 3-NT molecule, the main limitation of this
electrochemical sensing platform was that detection levels should be improved. This could be
solved by employing a more favourable biorecognition approach.
In order to enhance both selectivity and sensitivity of the previous approach, a molecular
imprinting technology for 3-NT is presented in Chapter 6. Herein, the incorporation of this
strategy with the previous wax-printed paper-based electrode enabled to create specific sites of
biorecognition towards the direct electrochemical detection of 3-NT molecule. In addition, the
applicability of this biosensor was tested and validated through assays performed in urine
samples.
Finally, it was important that a multi-analyte detection could be achieved to solve current POC
needs. This could be achieved by having one electrode per analyte or trying to have a single
electrode that enables a multi-analyte detection. This last option was selected, for being
simpler. Thus, Chapter 7 focused on the fabrication and characterization of core-shell magnetic
nanoparticles modified with a redox-active probe. The use of magnetic nanoparticles meant a
pre-concentration effect at the sensing surface and the modification of the redox-active probe
signalled which biomarker was being targeted. Under this concept, a specific redox probe binds
to a specific OS biomarker bound to a molecularly-imprinted support, and the electrical signal of
the bound probe should identify the biomarker involved and its quantity. For this purpose, each
redox probe to be read should be linked to a different OS biomarker. This concept was tested
herein for a single redox probe, but it is important to highlight that it could be further extended to
other compounds. First, the experimental conditions during the synthesis of the nanoparticles
were studied and optimized, in order to tune the size and electrochemical performance of the
Chapter 1
4
material. Next, the redox probe also needed special tuning to ensure that its electrochemical
signal was the only signal generated at its redox potential range.
Finally, Chapter 8 summarizes the main conclusions of this project and highlights potential
applications for future research work.
1.4 LIST OF PUBLICATIONS
1.4.1 Papers published in international scientific journals
- Gabriela V. Martins, Ana C. Marques, Elvira Fortunato, M. Goreti F. Sales, "Electrochemical
paper-based biosensor for label-free detection of 3-nitrotyrosine in human urine samples using
molecular imprinted polymer", submitted.
- Gabriela V. Martins, Ana C. Marques, Elvira Fortunato, M. Goreti F. Sales, "Wax-printed
paper-based device for direct electrochemical detection of 3-nitrotyrosine", Electrochimica Acta
(2018), 284, 60-68.
- Gabriela V. Martins, Ana P. M. Tavares, Elvira Fortunato, M. Goreti F. Sales, "Paper-Based
Sensing Device for Electrochemical Detection of Oxidative Stress Biomarker 8-Hydroxy-2′-
deoxyguanosine (8-OHdG) in Point-of-Care", Scientific Reports (2017), 7, 14878-14887.
- Gabriela V. Martins, Ana C. Marques, Elvira Fortunato, M. Goreti F. Sales, "8-hydroxy-2′-
deoxyguanosine (8-OHdG) biomarker detection down to picoMolar level on a plastic antibody
film", Biosensors and Bioelectronics (2016), 86, 225-234.
1.4.2 Communications presented in national and international scientific
conferences
- Gabriela V. Martins, Stefano Chiussi, Goreti F. Sales, "Fabrication of flexible sensing devices
for application in cancer diagnosis", IV Jornada Científica de IBEROS, Lugo (Spain), 11 July
2018 (oral presentation)
- Gabriela V. Martins, Ana C. Marques, Elvira Fortunato, M. Goreti F. Sales, "Electrochemical
paper-based sensor integrated with molecular imprinting towards point-of-care diagnosis", 6th
World Congress and Expo on Nanotechnology and Material Science, Valencia (Spain), 16-18
April 2018 (oral presentation).
- Gabriela V. Martins, Ana C. Marques, Elvira Fortunato, M. Goreti F. Sales, "An imprinted
paper-based biosensor designed towards point-of-care diagnosis", Graduate Student
Symposium on Molecular Imprinting, Porto (Portugal), 8-9 June 2018 (oral presentation).
Chapter 1
5
- Gabriela V. Martins, Ana C. Marques, Elvira Fortunato, Helena R. Fernandes, M. Goreti F.
Sales, "Carbon-based electrodes on paper substrates for biosensing purposes", iBEM -
International Biomedical Engineering Meeting, Porto (Portugal), 21 March 2018 (poster
presentation).
- Gabriela V. Martins, Elvira Fortunato, Helena R. Fernandes, M. Goreti F. Sales, "Chip-on-
Paper for sensoring 8-hydroxy-2'-deoxyguanosine (8-OHdG) oxidative stress biomarker in point-
of-care", NanoPT2016 International Conference, Braga (Portugal), 16-19 February 2016 (oral
presentation).
- Gabriela V. Martins, Elvira Fortunato, Helena R. Fernandes, M. Goreti F. Sales, "A molecularly
imprinted sensor for sensitive detection of 8-hydroxy-2'-deoxyguanosine (8-OHdG) oxidative
stress biomarker", Graduate Student Symposium on Molecular Imprinting, Kent (UK), 27-28
August 2015 (poster presentation).
- Gabriela V. Martins, Elvira Fortunato, Helena R. Fernandes3, M. Goreti F. Sales, "A
biomimetic sensor for monitoring oxidative stress biomarker in point-of-care", 1st ASPIC
International Congress, Lisbon (Portugal), 25-26 November 2014 (poster presentation).
Chapter 1
6
Chapter 2
7
CHAPTER 2
2 Introduction
Along this chapter, a brief background concerning the fundamental principles of
biosensors, the importance of biomarkers and their application in the context biomedical
devices is presented and discussed. In addition, the integration of novel electrode systems
based on modified-conductive substrates is also explored, as a potential alternative in the
design of low-cost, efficient and portable sensor platforms.
2.1 OXIDATIVE STRESS
An overview of the current literature highlights the diversity and, in some cases, divergent
theories proposed to understand the aging process and their mean features. Generally, "aging"
can be defined as the sum of all the mechanisms that can, direct or indirectly, modify the
functions of a living being, by preventing it from maintaining physiological balance and,
consequently, causing the death of the organism. One of the first theories proposed by Harman
in the 1950s was the so-called "free radical theory" claiming that the production of free radicals
among aerobic organisms was determinant for causing cellular damage [5]. Under this scope,
the definition of "free radicals" comprises atoms and molecules composed by unpaired and
highly reactive electrons in their outer orbits, making these radicals quite unstable and highly
reactive [6].
Afterwards, it was found that reactive oxygen species (ROS) greatly contributed to the
accumulation of oxidative damage to cell macromolecules, but this kind of molecules were not
exclusively free radical species. So, ROS are metabolites of molecular oxygen (O2) holding
higher reactive behaviour than O2 that includes superoxide radical (O2-), hydroxyl radical (OH
-)
and hydrogen peroxide (H2O2), among others. Interestingly, in normal physiological conditions,
ROS are continuously generated as by-product species during normal aerobic phenomena, as a
response to stress stimuli, pathological conditions or even to external environment effects [3].
Furthermore, as a protective barrier against these harmful compounds, one can find a complex
system of antioxidant and protecting species (see Figure 2.1), such as, superoxide dismutase,
glutathione peroxidase, glutathione reductase, catalase, glutathione and vitamins C and D [7].
Chapter 2
8
Figure 2.1: Adapted scheme highlighting the various activators and inhibitors factors associated to the
production of ROS [7].
Nowadays, some believe that our cell metabolism may be the source of our aging process. This
"oxidative stress theory" claims that high levels of ROS are directly linked to damage of
macromolecules, like proteins, nucleic acids and lipids [8]. Although low levels of ROS enable
cell signal transduction and immune response, a high accumulation of ROS in the organism can
overload cellular macromolecules under OS circumstances, resulting in some human diseases,
specially from cardiovascular and neurodegenerative origin. These include, Alzheimer,
Parkinson and Huntington disease, and also some cancer pathologies [9][10]. In this regard,
recent studies have shown that high concentrations of biological markers originated from OS
occurrence can be associated with degenerative diseases, hypertension, type II diabetes and
several types of cancer [11]. In parallel with ROS, reactive nitrogen species (RNS) can also
occur, enabling analogous nitrosative effects.
Nevertheless, ROS also play a crucial role as signalling molecules in response to changes in
intra- and extracellular environment, aiming to ensure the maintenance of physiological
functions [7]. Currently, this is a controversial issue around ROS, besides their active part in the
damage of relevant biomolecules, studies along the years have also confirmed their importance
through the mechanisms involved in the antioxidant defence systems [12]. In addition, at the
cellular level, ROS can be originated from both endogenous and exogenous sources [13][14],
as seen in Figure 2.2. Although ROS produced by exogenous processes that includes, ozone,
ultra-violet radiation and other environmental pollutants can cause OS, in most circumstances,
the endogenous ROS-originated the main threat responsible for attacking cell molecules.
Importantly, not all endogenous ROS production hold negative consequences, occurring some
Chapter 2
9
controlled and relevant enzymatic reactions that are important for cellular maintenance function
[3]. In particular, the endogenous antioxidant defences include a network of antioxidant
enzymes, such as, superoxide dismutase, gluthathione peroxidase and catalase that are widely
distributed in the cells holding the ability to counterbalance oxidative stimulus. Thus, on a
normal and continuous basis, all living organism are exposed to ROS, in aerobic environment
conditions.
Figure 2.2: Adapted scheme showing the proposed mechanisms for ROS production and their
contribution to the aging process [14].
Most cancer pathologies can be commonly defined by three main stages that are: initiation,
promotion and progression [15]. Along all phases, OS plays a relevant role, specifically, during
cancer initiation and progression. ROS can increase DNA mutations and also contribute to a
growth in cell proliferation or a decrease in apoptosis of the initiated cell population. In addition,
OS will also participate in the progression stage by introducing further DNA alterations to the
initial cellular population [7]. Overall, a close but still not clear link has been commonly proposed
by epidemiological and experimental data between the development and progression of cancer
pathologies and OS phenomena. In this context, the physiological level of relevant biological
markers presented in various biological tissues or fluids can be also influenced by external
conditions, such as, age [16] and gender [17], for instance. Thereby, the importance of
Chapter 2
10
establishing and understanding early biological markers has been a growing need in order to
accomplish novel therapeutic approaches and, consequently, reduce disease's mortality.
2.1.1 Biomarkers of Oxidative Stress
In order to assess OS in biological matrices, two different approaches can be employed:
(i) the increase of the levels of ROS can be directly quantified or;
(ii) the damage resulted from OS can be measured.
There are some studies related to the direct detection and quantification of ROS in complex
biological samples but their main limitations resulted from their low stability, high reactivity and
short half-lives [18][19]. Therefore, the most common way is not measuring the level of ROS
themselves, but quantify the damage originated from these, since in terms of consequences to
the organism, it is the outcome of OS that matters rather than the total amount of ROS
originated. As mentioned before, the radical species that are formed during OS phenomena will
interact directly with the biomolecules present in the cells, such as, proteins, phospholipids and
nucleic acids, causing cell degeneration and death [13][20]. In parallel, specific molecules are
produced and their quantification can be used as OS biomarkers for different biological matrices
[21]. Overall, biomarkers of OS can be designated as molecules that are modified by suffering
interactions with ROS in a biological microenvironment.
According to IUPAC recommendations, "biomarker" is defined as a parameter that can be used
to identify a toxic effect in an individual organism and can be applied in extrapolation between
species, or as an indicator signalling an event or condition in a biological system or sample and
giving a measure of exposure, effect, or susceptibility [22]. Moreover, a successful biomarker
should have (1) high specificity for the effect of interest; (2) high sensitivity; (3) reflection of an
early effect; (4) easy and inexpensive analysis; (5) low background level of the biomarker in the
body fluid of interest; (6) a well-established relationship between the response of the biomarker
and exposure, and (7) a well-established relationship between the response of the biomarker
and the induced damage [10].
Under this scope, a relevant and global biological indicator of oxidative damage is not yet
available, mainly because measuring OS is quite complex and a combination of different
measurements may be required. Although the key issue to the assessment of oxidative status
of an organism may be the analysis and quantification of different OS biomarkers, special care
must be taken during the interpretation of the results [12]. Thus, the discovery and selection of
robust biomarkers is crucial for screening, early diagnosis, and monitoring relevant diseases.
Case-studies have investigated a possible correlation between the impact of environmental
factors, like air pollution or smoke, and the occurrence of OS but results were not conclusive
[23]. Moreover, the identification and validation of potential biomarkers may be hindered by their
low levels in the biological fluids and the intrinsic variability among control and patient samples.
Several studies related to neurodegenerative disease diagnostics and monitoring have been
discussed in order to conclude that, by analysing the inflammatory profile in association with
Chapter 2
11
disease-specific biomarkers, the diagnostic and prognostic could potentially be improved [24].
Thus, there is a need to work with a wide screen of possible biological markers. According to
the origin of the biomolecule, OS biomarkers can be categorized in the following way (see Table
2.1):
Table 2.1: Summary of the most relevant OS biomarkers and their main categories.
MEASURE Biological Marker Reference
DNA oxidation
8-Hydroxy-2'-deoxyguanosine [25–30]
8-Hydroxyguanosine [31,32]
Lipid peroxidation
Malondialdehyde [33–38]
4-Hydroxynonenal [39,40]
8-Isoprostane [41]
Acrolein [42,43]
Protein oxidation
3-Nitrotyrosine [44–48]
3-Chlorotyrosine [49]
2-Pyrrolidone [50]
Although most of the performed studies have assessed the presence of OS biomarkers in
blood, serum, plasma or urine samples, other biological fluids such as, cerebrospinal fluid, nasal
lavage fluid, joint fluid, breast milk, tissues or exhaled breath have been successfully used. So,
the application of different analytical methodologies to a wide diversity of biological samples
have improved the accuracy of the biomarker assays. Meanwhile, great efforts are still on-going
in order to develop a suitable screening program targeted for POC assessment. Beyond
prevention, early detection is the most crucial determinant for successful treatment and survival.
So far, current methods for cancer diagnosis based upon pathological examination alone are
insufficient for detecting early tumour progression. Under this scope, the potential of using OS
biomarker assays for early detection purpose has been acknowledged as a determinant tool for
successful cancer treatment, monitoring and survival.
Chapter 2
12
2.1.1.1 8-Hydroxy-2'-deoxyguanosine
The most commonly studied biological markers of DNA damage obtained through the
attack of nucleotides bases are 8-OHdG and 8-hydroxyguanosine (8-OHG). Both 8-OHdG and
8-OHG products constitute oxidation derivates of guanine, one of the four main nucleobases
found in the nucleic acids DNA and RNA. In this context, the formation of 8-OHdG through DNA
hydroxylation is an important mechanism of oxygen-radical induced mutagenesis [51] and has
already been acknowledged as a suitable biomarker of OS [52]. For now, high-performance
liquid chromatography (HPLC) has been the popular choice for determining the levels of this
kind of biomarker. In addition, 8-OHdG is the most widely used fingerprint of DNA damage,
yielding strong implications along carcinogenesis evolution. Two possible mechanisms can be
presented concerning the formation of 8-OHdG, as can be seen in Figure 2.3:
Figure 2.3: Adaptation of the proposed mechanism for the formation of 8-OHdG [16].
Different methodologies have been proposed to analyze 8-OHdG in biological samples,
including Resonance Rayleigh Scattering (RRS) [53], Gas Chromatography - Mass
Spectrometry (GC-MS) [54], Enzyme-Linked Immunosorbant assay (ELISA) [55] and HPLC
[56]. Although these conventional detection techniques have reached good selectivity and
suitable detection limits, they hold strong limitations, such as, time consuming and often require
the use of sophisticated and laborious technologies, highly qualified personnel, and excessive
Chapter 2
13
handling of biological samples. Specifically, biochemical methods such as ELISA can undertake
cross-reactions, which can give false-negative or false-positive results. Furthermore, in most
cases, GC-based methodologies operate with pre-treatment procedures prior to qualitative and
quantitative analysis, originating long analysis periods. So, the development of fast, sensitive,
easy-to-use and low cost approaches for 8-OHdG detection remains a continuous challenge.
Over the last years, 8-OHdG biomolecule has been quantified in various biological samples,
such as, urine [57], saliva [58], blood [29] and tissue [59]. It was found that the average levels of
8-OHdG in healthy humans is around 20 ng/mL, a value that increases when OS rises [60].
Bolner et al. have demonstrated that 8-OHdG levels in Parkinson's disease patients can be 2-3
times higher than in healthy controls [29]. Therefore, highly sensitive (nanomolar level)
methodologies are needed for the assessment of 8-OHdG.
Although the detection of oxidative damage is being widely applied in human research and
clinical applications, we still need more data about the main factors that determine the basal
levels of these biomarkers among general population. Meanwhile, some interesting studies
have suggested that some variables, such as, age, sex, alcohol consumption, level of
education, the time season of sample collection and exposure to heavy metals are implicated
with 8-OHdG quantification [61][62]. Thus, special care must be taken during data analysis
because the occurrence of this type of OS biomarker can be monitored not only in body fluids
and tissues of patients, but also in healthy people in physiological (and variable) concentrations.
2.1.1.2 3-Nitrotyrosine
Looking more closely into the OS pathway, one finds nitric oxide (NO), a reactive nitrogen
specie that plays an important role in many pathologies, such as, ischemia-reperfusion, septic
shock, neurodegenerative and chronic inflammatory diseases [63]. The overproduction of this
nitric oxide can generate a cytotoxic molecule called peroxynitrite (ONOO˙) that is responsible
for attacking protein residues [64]. As a consequence, 3-NT is obtained as a stable final-product
and has been indicated as a biomarker for OS diagnosis [63]. Although peroxynitrite has been
regularly correlated with increased oxidative reactions and DNA damage in inflamed tissues
[63][64], it is not the only source of production of 3-NT marker.
Free 3-NT can be synthesized in-vivo by the direct reaction of tyrosine with nitrating oxidants
(see Figure 2.4), which includes mainly three distinct biochemical processes [44]. Briefly, (i) the
superoxide anion (O2˙-) is the product of the one-electron reduction of oxygen being the
standard reduction potential for this conversion from molecular highly dependent on the nature
of the medium; (ii) nitric oxide is an highly reactive free radical that is formed in mammalian cells
during the oxidation of L-arginine to L-citrulline•−
; and finally (iii) peroxynitrite that, contrary to
nitric oxide and superoxide anion is not a free radical specie, but constitutes a strong oxidant
and a nitrating agent that can originate the nitration of tyrosine residues in proteins. Thus,
tyrosine nitration is the permanent addition of a nitronium group (+NO2) at the ortho-position
Chapter 2
14
resulting in free or protein-associated 3-NT. Consequently, the formation of nitro-aromatic
compounds such as 3-NT can be directly associated with nitrosative stress [65].
Figure 2.4: Adapted scheme showing the different pathways responsible for the nitration of tyrosine
residues of proteins and consequent production of 3-NT [44].
Meanwhile, independently of the route of formation of 3-NT (via oxidation of nitric oxide, via
peroxynitrite formation, etc.), the quantitative estimation of 3-NT, either free or bound to
proteins, has been employed as a marker of oxidative damage in chronic inflammation [66],
cardiovascular [67], atherosclerotic [68] and tobacco smoke–associated lung diseases [69].
Moreover, as mentioned before, the reactive species responsible for leading to tyrosine nitration
usually have very short half-lives and, consequently, their direct measurement in biological
environment is highly difficult.
In this context, 3-NT has been detected in several biological tissues and fluids, such as, plasma,
serum, urine, cerebrospinal fluid, tissue samples, among others. Aiming for an accurate
quantification, different methodologies have been proposed including immunoassays [48], GC-
MS [70][71], liquid chromatography methods coupled with ultraviolet [72][73], fluorescence [74],
mass spectrometry [65] and electrochemical detection [75]. Therefore, the application of the
above techniques enabled very low limits of detection with high accuracy standards, but still
involved complex and time-consuming sample preparation steps that are not compatible with
routine in-loco analysis. Moreover, current detection methods are quite expensive and
impractical for scaling up [76]. Since 3-NT was suggested as a biomarker of OS, a growing
Chapter 2
15
effort has been made to implement portable, facile and rapid analysis of clinical samples in POC
screening, as a way to improve the outcome of prevention and therapeutic approaches.
2.1.1.3 Malondialdeyde and Hydroxynonenal
As mentioned earlier, lipids are one of the major targets of oxidative attack and, the
modification of these molecules can increase the risk of some degenerative diseases [77]. Once
lipid peroxidation is initiated, the propagation of chain reactions will occur until a number of
secondary, but highly damaging products are produced. Among others, the two most used end
products of lipid peroxidation are MDA and 4-HNE (see Figure 2.5).
Figure 2.5: Chemical structure of A) malondialdehyde and B) 4-hydroxynonenal.
One of the most common techniques to quantify MDA levels is through the spectrophotometric
thiobarbituric reactive specie (TBARS) assay. Although this methodology is quite easy and
straightforward to be implemented, it is a non-specific test that can detect aldehydes other than
MDA, with some artefact issues related to sample preparation [76]. Moreover, MDA data
obtained through TBARS approach seem to be highly influenced by the smoking status,
meaning that a wide variety of studies have found a significantly increased level of MDA in
smokers in comparison with non‐smokers [76].
MDA quantification as an OS biomarker has been used in various biological matrices, such as,
saliva [37], urine [77], plasma [78], tissue [38], among others. Another technique also employed
to assess MDA levels in biological samples is by using an HPLC approach [79]. Therein, a
comparative study against the traditional TBARS assay have showed an overestimation of the
MDA concentration found during the TBARS methodology. Nevertheless, MDA is still employed
as a known biomarker of OS in various pathologies despite its basal level variations among
healthy people [80][81][82].
The occurrence of 4-HNE, a sub-product originated as a consequence of oxidative damage, has
been commonly associated to onset of memory loss and cognitive dysfunction, characteristic
symptoms of Alzheimer’s disease [83]. Furthermore, 4-HNE is a potent modulator of numerous
Chapter 2
16
cell processes and can be accumulated during numerous oxidative stress-related diseases,
such as neurodegenerative and cardiovascular diseases, metabolic syndrome and also cancer
[84][85]. Some studies have performed 4-HNE quantification by using an enzyme immunoassay
approach [39], fluorescence immunoassay [86] liquid chromatography tandem mass
spectrometry (LC-MS) [87], TBARS assay [88]. Interestingly, 4-HNE should not be found in
urine fluids so, urinary biomarkers for 4-HNE offer a non-invasive biomarker of lipid peroxidation
and OS [89]. In sum, one still do not reach a consensus related to the cut-off of some individual
biomarkers of OS, but their important relevance in the genesis of many diseases have been well
established, making them ideal candidates to be investigated and applied in early screening.
2.2 BIOSENSORS
In the last decades, along with the digital and technological evolution, biosensor research
has boosted in different direction areas, such as, healthcare, environment, food safety, industry,
pharmacology, sportswear, regenerative medicine, among others [90][91][92][93][94]. In
general, a biosensor is a chemical sensing device comprising a biologically derived recognition
element coupled to a physicochemical transducer [95][96]. Mainly, biosensor devices can
incorporate three distinct parts: (i) the biological component that recognizes the target molecule
resulting in a signalling, (ii) the transducer platform and, finally, (iii) a reading equipment
responsible for the data output (see Figure 2.6).
Figure 2.6: Schematic representation illustrating the main constitution, function and practical application of
a biosensor device.
Chapter 2
17
2.2.1 Recognition Element
Molecular recognition has been the crucial phenomenon for biosensing, as it is
responsible for distinguishing the target analyte from many other analytes present in a sample.
Chemical sensors and biosensors may be classified according to their type of biorecognition
element, being the most common ones, enzymes, nucleic acids, antibodies and living
organisms, such as, cells and cell organelles [97]. Among these, one can also include two
distinct classes of sensing elements, namely, the catalytic sensors that are related to enzymes,
microorganisms and other biomimetic catalysts, and the affinity-based sensors that incorporate
nucleic acids, antibodies and synthetic receptors [98].
Specifically, catalytic enzyme based-biosensors have become a prominent and widely chosen
approach making use of specific reaction products that can be easily measurable, like, electrons
and protons. The main advantages of using enzymes as recognition elements are also due to
their high selectivity and sensitivity response. In contrast, these catalytic enzymatic sensors can
be quite expensive and, in some cases, during immobilization process the catalytic activity may
become compromised.
Meanwhile, affinity biosensors imply the immobilization of a specific recognition layer, such as,
antibodies, nucleic acids and other affinity proteins, on the surface of the electrode in order to
selectively recognize the target biomolecule. Thus, the application of antibodies as the
recognition element is among the most popular choices for quick detection purposes, due to the
high sensitivity and specificity of antibody-antigen interaction [99]. Despite this popularity,
antibodies hold some limitations associated to their solubility and stability in aqueous conditions.
Concerning typical nucleic acid-based sensors, the biorecognition process involves non-
covalent interactions between the complementary bases. Moreover, as recognition elements,
nucleic acids are chemically more stable in comparison with their counterparts antibodies.
In parallel, the design of the sensor can also be differentiated by way the bio-receptor is
immobilized that includes, physical sorption, covalent binding, immobilization in a polymer
matrix, covalent binding and affinity immobilization [100]. This constitute an important and
crucial step that will affect the performance and efficiency of the sensor device. Generally, the
adsorption approach is the most simple and easiest to use, but the stability of the binding is
limited. In contrast, covalent interactions hold a longer lifetime due to the stronger bond
formation between the biomolecule and the solid support.
2.2.2 Signal Transduction
Another way to categorize biosensor devices is based on their transduction mechanism
(see Figure 2.7). In this regard, one can have three main types of transducers, which are the
piezoelectric (Quartz Crystal Microbalance/ QCM) [101–103], the optical (Colorimetric,
Fluorescence, Surface Plasmon Resonance/ SPR) [104,105] and the electrochemical [106–108]
biosensors.
Chapter 2
18
Figure 2.7: Schematic representation with the main features of the different types of sensor transduction.
Piezoelectric sensors are mass-sensitive sensors that, by tracking the resonant frequency of
quartz crystal, are capable of measuring mass variations on the surface of the quartz electrode
[109]. Optical sensors use light as the way of transducing and their main advantage is the quick
and facile ability to give a visual response of the presence of a target molecule [110]. Among
these sensors, a wide variety of optical contrasts, such as, absorption, fluorescence, Surface-
Enhanced Raman Scattering (SERS) and refraction are used to detect optical changes. And
finally, an electrochemical sensor holds an electrochemical transducer, measuring the variation
of electrical current, potential, conductance or impedance at the interface of the electrode-
sample [111]. In addition, other physical properties, such as, heat and magnetic field can be
also employed for the conversion of the biological signal during the design of calorimetric and
magnetic based-sensors, respectively.
In sum, the continuous development and application of these novel biosensor devices is the
result of the combination of various advantages, such as [94]:
(i) high sensitivity and affinity toward their target molecules;
(ii) a suitable biological recognition that enables very selective sensing materials;
(iii) good and proper limits of detection targeted to the analyte of interest;
(iv) and their ability to be easily miniaturized in a ready-to-use apparatus;
Chapter 2
19
So far, biosensing approach constitutes a suitable choice to be integrated in a POC device
enabling special features, like, portability, automation and quick in-loco analysis. Among all
transduction schemes, electrochemical-based biosensors are currently the most promising
approach, where cost and minimized size are crucial needs.
2.3 ELECTROCHEMICAL BIOSENSORS
To overcome the growing need to develop biosensors with higher sensitivity and
selectivity, electrochemical detection has been acknowledged as the more suitable strategy
[112]. Electrochemical sensing technology has proven to be a valuable candidate for future
clinical diagnosis, targeting relevant biomolecules, such as, proteins, nucleic acids and other
small molecules [113]. Recently, an interesting review has compiled and discussed the
development of versatile electrochemical genosensing of circulating biomarkers associated with
cancer, neurodegenerative diseases and viral/ bacterial infections [114].
In general, an electrochemical sensor is designed to transform the effect of the physicochemical
interaction between analyte and the electrode surface into an electrical signal [115]. It has been
a very promising approach, due to its advantages, like high sensitivity, selectivity, a broad linear
range of application and low cost instrumentation, enabling to overcome the most common
challenges of analytical analysis.
Currently, a wide diversity of electrochemical studies have been conducted with biosensors,
including the development of enzyme-based detection systems [116]; the establishment of
microfluidic platform for therapeutic drug monitoring [117]; the construction of carbon-based
paste for antioxidant estimation in wine samples [118]; the preparation of smart conductive films
for cell culturing [119]; or the implementation of multiplexed immunoassay targeted for cancer
biomarkers [120], among others. The use of electrochemical biosensors can also be performed
under two different ways of transduction, direct or indirect, if a redox mediator is required or not
to promote reversible electrochemical processes, enhancing the sensitivity and improving the
detection limits. Moreover, according to the type of transducer, electrochemical techniques can
be organized into conductometric, potentiometric or voltammetric biosensors or electrochemical
impedance spectroscopy (EIS) biosensors [96].
In voltammetric biosensores, a potential is applied to a working electrode versus a reference
electrode and the measured current results from an electrochemical oxidation or reduction of
the electroactive species. Usually, the transducer surface is a metal or carbon electrode that
can be chemically modified. Voltammetric biosensors allow to directly correlate the resulting
current measured at a constant potential value with the bulk concentration of the analyte specie.
Conductometric biosensors measure changes related to the electrical conductivity of the sample
solution during a chemical reaction, generally by using interdigitated microelectrodes. One of
the typical advantages of working with these type of biosensors is that their capacitance
measurements constitute a good indication of the insulating properties of the system.
Chapter 2
20
In potentiometric biosensors the potential difference of an electrochemical cell is measure, in
negligible (near-zero) current conditions. Usually, the transducer consists of an ion-selective
permeable membrane, being pH electrodes the most common potentiometric devices among
analytical laboratories.
EIS biosensors monitor the response of an electrochemical system (cell) to an applied potential,
in which the frequency dependence may reveal underlying chemical processes. It indicates the
resistive and capacitive properties of materials when a low amplitude sinusoidal perturbation is
applied [121]. EIS is routinely used for the characterization of functionalized electrode surfaces,
enabling sensitive measurements related to surface phenomena or variation of bulk properties.
Overall, the electrochemically-based biosensors that are most explored in recent scientific
papers for POC purposes are the ones of voltammetric and EIS nature. These are addressed
next in more detail.
2.3.1 Voltammetry
As mentioned before, voltammetric based-biosensors allow tracking the concentration of
the target-analyte as a function of current variation. Typically, voltammetry is characterized by
the occurrence of an oxidation or/ and a reduction reaction of an electroactive(s) specie(s) at
the electrode surface. Briefly, the current will change accordingly to the kinetic and mass
transport involved during the reaction and diffusion of the species [122]. As an alternative to the
conventional approaches, some biosensing devices targeted for OS biomarkers have been
reported in the literature, relying mostly on electrochemical or QCM detection modes. For
instance, such devices establish a direct reading of 8-OHdG, making use of its active redox
properties on glassy carbon [27][123] or pyrolytic graphitic [124] supports, modified with highly
conductive nanomaterials. In general, the detection capability of these devices varies from 1.1
to 97 nM, but some methods have shown severe interference from urine elements (mostly uric
acid), solved therein by the addition of specific enzyme prior to the analysis stage. Furthermore,
the strategy behind electrochemical biosensors designed as relevant tools of oxidative damage
allows to track directly their oxidant activity and simultaneously, their biological and
physicochemical features as clinically relevant biomolecules. Nevertheless, this kind of
electrochemical technique is known to enable the detection of low concentrations of analyte and
different voltammetric approaches can be used according to the way the potential is applied.
2.3.2 Cyclic Voltammetry
The most common voltammetric technique used for primary research concerning the
electrochemical behaviour of biomolecules is cyclic voltammetry (CV). CV measurements allow
to monitor current variation while the potential applied on the working electrode is reversibly
varied, under a fixed scan-rate value. As can be seen in Figure 2.8, cyclic voltammogram is the
graphical representation of the current response as a function of the applied potential and the
most relevant parameters to evaluate a redox system include:
(1) Ipa and Ipc, as the anodic and cathodic peak currents, respectively;
Chapter 2
21
(2) ʋ, as the scan-rate of the applied potential;
(3) Epa and Epc, as the anodic and cathodic peak potentials, respectively;
Figure 2.8: Cyclic voltammogram obtained for a reversible system.
The reversibility of a redox process is related to the rate at which the electron transfer occurs
between the electroactive species and the electrode surface [125]. Firstly, in a reversible
electrochemical process the Nernstian equilibrium is rapidly attained. Moreover, CV is a widely
used tool to assess the reversibility of the redox reactions by simple measurement of
voltammograms at distint scan-rates. Thus, one of the requirements of reversible
electrochemical processes is a clear linearity between the peak current and the square root of
the scan-rate. In parallel, this outcome constitutes also a intrinsic characteristic of a diffusion
controlled process.
Another way to evaluate the electrochemical reversibility of a redox reaction is the difference
between the anodic and cathodic peak potentials (ΔEp = Epa - Epc). So, for a reversible ideal
electron transfer, Ep should be independent of ʋ and ΔEp equal to 59/n mV (at 25ºC), being n
the number of electrons involved in the electrochemical reaction [122]. Finally, the ratio between
the anodic and cathodic peak currents (Ipa/Ipc) is also employed to estimate the reversible
behaviour of the redox system. In this case, the electron transfer is reversible when this ratio is
close to 1 and for a less reversible system the ratio will move away from 1.
By contrast, in irreversible systems the linear potential sweep and CV lead to the same
voltammetry profile because no reverse peaks appear upon changing the sweep direction [122].
Under these conditions, concentrations of the oxidized and reduced species do not follow the
Nernst equation and the individual peaks are reduced in size and widely separated, as can be
seen in Figure 2.9.
Chapter 2
22
Figure 2.9: Comparison of typical voltammograms obtained for reversible and irreversible systems.
2.3.3 Square Wave Voltammetry
Square wave voltammetry (SWV) has been recognized as a crucial tool for sensible
detection of relevant biomolecules, such as, proteins, nucleic acids, etc. Generally, in SWV the
excitation signal is composed by a symmetrical square-wave pulse of amplitude undertaken
with a staircase waveform of step height, being the forward pulse of the waveform equal to the
staircase step [113]. In addition, the net current is determined by measuring the difference
between the forward and reverse currents. The main advantages of this voltammetric technique
are higher scan-rate during the analysis, less consumption of electroactive species and better
performance related to passivation of the electrode surface [122].
2.3.4 Differential Pulse Voltammetry
The use of methodologies based on small-amplitude pulses have enhanced the
sensitivity and accuracy parameters of electro-analysis. Under this scope, differential pulse
voltammetry (DPV) is a technique that enables the application of the potential with a series of
sequential pulses with fixed and continuous amplitudes. The current is measured immediately
before the pulse application and after the end of the pulse, being determined the difference
between both current values [122]. The peak representation plots the current difference for each
pulse point as a function of the potential.
Chapter 2
23
2.3.5 Electrochemical Impedance Spectroscopy
Electrochemical impedance spectroscopy (EIS) is another powerful electrochemical tool
with great interest for the detection of relevant chemical and biological species. It has the
advantage of not causing any damage or disturbance to the analysed surface of the
electrochemical system. Mathematically, impedance has been expressed as a complex number
comprising resistance (real part) and reactance (imaginary part) [126]. By looking to the Ohm’s
law (equation 2.1), where R is the resistance, V is the potential and I is the current:
(2.1)
whereas, the impedance of the system (Z) can be calculated by means of an expression
analogous to the previous Ohm's law, being Et the applied potential and It the resulting current:
(2.2)
The applied potential Et can be expressed as a function of time:
(2.3)
where E0 is the amplitude of the voltage signal at t = 0, ω is the angular frequency ( )
and t is time.
Also, the resulting current It has a phase shift θ with amplitude of I0, which can be expressed by:
(2.4)
So, the expression for Ohm' Law can be applied to calculate the impedance of the system as:
(2.5)
The impedance, Z, can be expressed in term of a magnitude of Z0 and a phase shift, θ.
Then, by applying Euler’s relationship (equation 2.6) given by:
(2.6)
where Ɵ is a real number and j is the imaginary unit.
Chapter 2
24
So, it is possible to express the impedance as a complex function, being the potential described
as:
(2.7)
and the current is expressed as:
(2.8)
Therefore, impedance can be represented as a complex number:
(2.9)
(2.10)
And finally, the impedance is now represented in the form of real part (Z0 cosθ) and imaginary
part (Z0 sinθ):
(2.11)
(2.12)
The most direct interpretation of EIS measurements is typically performed by fitting the
impedance data to an equivalent electrical circuit that gives a representation of the physical
process that occurs in the interface system. Therefore, any process occurring in an
electrochemical cell can be represented in terms of equivalent circuits composed by different
combinations of resistors (R), capacitors (C) and/or inductors (L). Briefly, the equivalent circuit
of Randle's electrochemical cell is mainly composed by the following components [127]:
(1) RΩ, the solution resistance, which is dependent on the ionic concentration, type of ions and
electrode area;
(2) Rct, the charge-transfer resistance, which is inversely proportional to the electron transfer;
(3) Cd, the double-layer capacitance;
(4) Zw, the Warburg impedance that can be used to assess effective diffusion coefficients;
These processes occur at the electrode-electrolyte interface comprising a fast mass-transfer
reaction, a slow-paced charge transfer reaction and, finally, a diffusion phenomenon at the
interface. Nyquist plots, also known as Cole-Cole plots, are one of the most popular ways to
evaluate impedance information due to the facile prediction of the circuit elements as well as it
allows the graphical extrapolation to obtain the RΩ parameter (see Figure 2.10).
Chapter 2
25
Figure 2.10: Nyquist plot for an electrochemical Faradaic system.
One of the limitations of this kind of plotting is that the information concerning the frequency is
dismissed so, in order to overcome this issue, another way to express impedance data is also
with a Bode diagram, by plotting absolute Z(ω) and phase angle θ(ω) in the frequency domain
(see Figure 2.11).
Figure 2.11: A typical example of a Bode plotting.
Typically, an electrochemical reaction can occur on the electrode surface by two distinct limiting
mechanisms, such as, kinetically controlled (charge transfer based) and diffusion controlled
(mass transfer based). The Nyquist plot for a Randle's cell is constituted, at higher frequency,
by a single time semi-circle curve due illustrating the charge transfer process and a diagonal
line of the diffusion process (Warburg impedance), at low frequency [126].
Moreover, EIS experiments can also be divided in two different types, Faradaic and non-
Faradaic processes [127]. Herein, the main focus was given to the Faradaic processes where a
redox marker specie is needed, and so particular attention is given to the Rct. Under these
Chapter 2
26
conditions, the working electrochemical cell must contain both reduced and oxidized forms of a
benchmark redox probe that will undergo redox reactions at the electrode surface. In sum, the
application of EIS technique has been widely open to different areas of expertise, such as,
corrosion of metals, electrochemical synthesis of materials, study of ions mobility in energy
storage and biosensor devices [128][129][130][131][132][133].
2.3.6 Electrode size, materials and supports
Electrochemical measurements require an electrochemical cell that combines 2- or 3-
electrodes in electrical contact through the same solution. The electrodes involved are always
reference electrode (RE) and working electrode (WE), and an additional counter or auxiliary
electrode (CE) may be present, especially when the cell resistance is relatively high. In the 3-
electrodes configuration, the potential of the WE is monitored against the RE potential, and the
current passes between the WE and the CE. The RE remains a reliable reference for potential
control because it has ideal nonpolarizability since no (or little) current passes through it.
Experimentally, the tip of the RE is placed as close as possible to the WE in order to minimize
the solution resistance. Today, there is a variety of materials and approaches that can be used
to prepare the electrodes, with special relevance to the WE, which is designed to become
selective to a given target analyte, and the exact cell design varies with the specific needs of a
given experiment.
When the amount of analyte is not a concern, ranging from least 10 mL or more, a conventional
cell may be used, employing electrodes of tubular shape that round 50.5 cm, or more,
depending on the specific size by which these are manufactured. However, with limited
quantities, it is difficult to handle the corresponding electrical measurements in a very small
spot. Although the individual electrodes may be microsized and properly aligned to handle low
sample volumes, limitations from current/potential measurements might arise because of the
solution resistance and the heterogeneity of the electrolyte solutions.
The need to reduce the sample volume in electrochemical sensing has been targeted by
suitable miniaturization of the electrode combination in use, all together in the same small spot.
Screen-printed electrodes (SPEs) have become today a popular version of such miniaturization.
Along with minimal volumes of sample, this specific design offers portability, low-cost of
fabrication, large-scale production, and in many times short response time [112]. Following this
trend, printing technology has been explored and improved, aiming to obtain the deposition of
several successive layers of conducting material layers on an insulating support. At this point,
the most common printing techniques used nowadays are inkjet and screen-printing [134]. The
choice of the more suitable technique to perform the build-up of the sensing platform depends
on the type of material used and also the type of modification required. Briefly, a typical SPE is
composed by an inert substrate material holding three different electrodes together [93]. A
typical configuration designed specifically for the purposes of this work is shown in Figure 2.12.
Chapter 2
27
Figure 2.12: Main constituents of a paper-based assembly with three-integrated electrodes.
The analytical performance of the SPEs is directly tuned by the material used in the WE, while
enabling a real-time field analysis. Thus, this electrode mostly defines the overall characteristics
of the whole device and in most cases it can be printed using carbon, gold or silver inks.
Various factors can tune the electrochemical performance of the SPEs, mainly the properties of
the ink, the roughness of the surface, the printing process, the curing and drying temperatures,
among others [135][136]. There are already several commercially available SPEs (see Figure
2.13), keeping important features such as, material diversity, facile operation, enhanced
sensitivity, in some cases.
Figure 2.13: Examples of the wide diversity of SPEs commercially available.
Yet, many commercial versions still present reproducibility issues. An interesting report
regarding the electrochemical performance of different commercial SPEs showed that their
behaviour varied with the type of electrode, making this a critical issue for reproducibility [137].
Furthermore, this integrated electrically conducting layer can be assembled on different
chemically inert insulating materials, such as, ceramic [138], plastic [139], silicon [140], glass
[141] and paper [142]. In addition, the choice of the material for the working electrode results
Chapter 2
28
not only from its electrochemical reactivity features but also from its effect upon the background
current [143].
In terms of electrode material, metal-based electrodes, like for instance, platinum (Pt), gold (Au),
tin-oxide (SnO2) and copper (Cu), have been frequently used as working electrodes for an
electrochemical detection of biological molecules [144][145]. This is also the case in the context
of OS. Gutiérrez et al. reported self-assembled monolayers on gold-modified electrodes with
dendrimers for 8-OHdG detection enabling limits of detection around 1 nM [26]. Moreover, an
electro-chemiluminescence immunosensor based on platinum electrode modified with carbon
quantum dots and gold/silica (Au/SiO2) core-shell nanoparticles was designed for a rapid and
selective detection of 8-OHdG for biological samples [146].
The most common material applied in the sensing area of WEs in SPEs today is however
carbon-based. Carbon is one of the more important elements in nature and it can be found with
different structures holding very distinct characteristics, such as, graphite, diamond and
amorphous carbon. Carbon-based nanostructures present a wide range of interesting features,
namely, good mechanical, electrical and thermal properties, holding an important and quite
diversified potential use in many nanotechnology applications, as seen in Figure 2.14.
Figure 2.14: Representation of the different field of applications concerning carbon-based materials.
For instance, an electrochemical sensing approach on carbon materials based on the electro-
activity of DNA bases has been employed to measure its biomolecular damage [147][148][149].
In this context, guanine and adenine were immobilized as DNA bases on carbon electrodes and
their direct electrochemical response measured, as an indirect measure of DNA damage
detection [150][151]. In parallel, other approaches aim the direct determination of OS
biomarkers. In this case, carbon-based electrodes were widely applied, mostly due to their
physical and electronic characteristics [152]. Li et al. performed some electrochemical studies
Chapter 2
29
by using conducting polymer poly(3-methylthiophene) (P3MT) modified glassy carbon (GC)
electrodes for an electrochemical detection of urinary 8-OHdG, enabling a limit of detection of
0.10 M [153]. Moreover, Langmaier et al. investigated the oxidation reaction of 8-OHG on GC,
and compared it to other electrodes, made with Pt, Au or SnO2 [154]. It was observed that the
rate of charge transfer reaction depends on the nature of the electrode material, following the
sequence GC>Pt, Au>>SnO2. More interestingly, these effects can be related to the density of
the active surface sites (GC) or to the degree of oxidation of the electrode surface (Pt, Au,
SnO2). The determination of 8-OHG was also successfully tested on a graphitic support with
suitable antibodies [155].
The nanostructural materials derived from carbon are also attracting much attention. This
includes mostly carbon nanotubes (CNTs) and graphene, which constitute a good example of a
simple structure constructed from sp2 hybridized carbon bonds that, due to their high surface
area, make them good model systems to be used in nanoscale analysis. CNTs can be
distinguished by the number of layers that make up their cylindrical walls: single-walled
nanotubes (SWNTs) and multi-walled nanotubes (MWNTs). In terms of applying graphene as
the electrode material, researchers have showed that graphite control experiments have a huge
impact in the electron transfer properties of the obtained electrodes, enabling different kinds of
applications [156]. Overall, carbon-based nanomaterials, such as, graphene [157], carbon-black
nanoparticles [158], MW-CNTs [159] are being widely used as integrated structures for sensing
platforms, due to their singular electronic properties and controllable chemical functionalization.
As a practical example, the incorporation of carbon materials in the composition of conductive
inks to be printed as WE can highly improve the sensitivity of the biosensing device, making it
ideal tool for the development of biomedical diagnostics.
Moreover, these metal- or carbon-based electrodes have to be fabricated on suitable supporting
substrate. In SPEs, there are many different materials used for this purpose, such as, silicon,
glass, ceramic, polyvinyl chloride (PVC), etc.. During the selection of the substrate properties,
mechanical stability, facile fabrication and low-cost are important goals for the design of an
efficient device. Under this scope, cellulosic paper holds a great importance as an alternative
substrate material due to their unique features, like for instance, biocompatibility, high flexibility
and porosity, cost effective, high surface area and it is recyclable [160][161]. Due to the
incredible versatility of paper usage, the selection of the paper's type is highly dependent on the
specific assay and the analytical needs. Thus, a set of properties have to be well evaluated in
order to achieve the best paper-based analytical device [162], such as:
(i) Chemistry, the grade and distribution of the cellulose fibres, the presence of lignin, the
degree of esterification and also the content of nitrogen can affect the deposition, adsorption
and flow of species;
(ii) Surface area, it will have a great impact in the loading of the reagents, and also affects the
reproducibility and sensitivity of the assay;
(iii) Flow-rate, it represents the speed of a migrating specie along the paper path, affecting the
sensitivity of the paper-based device;
Chapter 2
30
(iv) Size of pores, as expected is related to the size of the particles that can be retained on the
paper surface;
(v) Porosity, it represents the void volume of the paper;
(vi) Thickness, it will dictate the speed of the analyte towards the testing area, meaning that thin
papers will be faster in comparison with thicker ones;
(vii) Cost, sustainability is a very important issue.
As the most abundant biopolymer in nature, cellulose paper has been employed as a support
material in analytical assays for centuries. The hydroxyl groups of the chains present in paper
are responsible for their hydrophilic behaviour making it an ideal choice for microfluidic systems
[163] in addition, studies have been performed related to how hydrophilicity of paper can be
finely tuned by modification with plasma treatments [164]. Over the years, paper has been
widely applied in different areas and the result was the development of new methodologies to
fabricate and modify paper devices. Usually, the goal of these methods is to produce
hydrophobic areas or channels on an hydrophilic porous paper substrate, being the most
common fabrication techniques wax printing, photolithography, wet etching, screen and inkjet
printing, among others, as seen in Figure 2.15 [165].
Figure 2.15: Adapted scheme with the various routes available to pattern paper-based sensors, with
respective main advantages and limitations [165].
Due to their simplicity, portability and low cost, the most commonly used analysis techniques
applied on paper-supported sensors are colorimetric [166][167], electrochemical [168][169],
chemiluminescent [170] and electrochemiluminescent [171][172] (see Figure 2.16).
Chapter 2
31
Figure 2.16: Schematic representation about techniques of nanofabrication, detection methodologies and
practical applications of paper sensors.
The most widely and cost-effective used technique has been the one giving a visual colour
alteration, but in most cases the results are only qualitative or semi-quantitative. Therefore, with
a high sensitivity and a good selectivity, electrochemistry appears as the second most common
transduction system. In addition, these paper-based biosensors are also aiming to become an
environmentally safe alternative to the conventional SPEs for screening OS biomarkers.
Moreover, in some cases, the reproducibility and sensitivity characteristics of the electrodes are
poor and/or the fabrication process is complex, thereby hindering scaling-up processes and
POC use. Table 2.2 illustrates some comparisons regarding the main features of the different
materials used as substrates in the design of sensor devices.
Table 2.2: Comparison of paper as a substrate material with other traditional materials [173].
Chapter 2
32
Another interesting benefit of paper over the traditional device materials includes portability,
affordability, user-friendly and small environmental impact, when compared to other supports,
making it a promising technology to be used for POC diagnostics. The high potential of using
paper as a substrate in biomolecule detection has increased exponentially in the last years, for
improving specificity and sensibility features. A quick overview on the last decade (see Figure
2.17) has shows the development of handmade paper-based immunosensor for electrochemical
detection of influenza virus [174]; paper-based colorimetric sensor toward bacteria detection
[166]; paper-based electrochemical cyto-device for sensitive cancerous cell detection [169]; and
also, a microfluidic immunoarray for simultaneous detection of cancer biomarker proteins [172].
In sum, the incorporation of this light and widely available natural resource has boosted the
development of compact and miniaturized electrochemical devices, holding high sensitivity and
selectivity, with great expectations towards early detection biosensing devices.
Figure 2.17: Some examples of different paper-based biosensors for A) virus [174], B) bacteria [166], C)
cell [169] and D) multi-protein [172] detection.
Chapter 2
33
2.4 MOLECULAR IMPRINTING POLYMER
Molecularly-imprinted polymer (MIP) has arrived to the new century has a highly
improved technology to update the conventional bio-recognition methodologies. Although this
concept is known since the early 1970s, this trend suffer a great income after the introduction of
a general non-covalent approach by Mosbach group [175]. In a simple way, the concept of
molecular imprinting is to design synthetic materials with the ability to mimic the behaviour of
natural biomolecules, including antibodies and enzymes, with the great advantage of being
highly stable and holding singular mechanical properties [176]. Thus, ideally, the capability of
imprinted materials to recognize the target molecule should be comparable to natural bio-
recognition in terms of sensitivity, affinity and selectivity. The assembly of MIPs (see Figure
2.18) can be divided into three fundamental steps: first, the template molecule to be imprinted is
mixed in a solution containing the block monomers, afterwards, through a polymerization
reaction the monomers (and crosslinker species) are "moulded" around the template creating a
polymeric matrix and, finally, the template molecule is removed from the matrix, creating specific
cavities with complementary size and shape [177].
Figure 2.18: Schematic representation of the synthesis of molecularly imprinted polymers.
Many considerations have to be made in order to achieve the optimal conditions for the
molecular imprinting process, such as, the choice of a suitable and compatible solvent for the
polymerization reaction; the need or not to introduce a cross-linking element to ensure the best
mechanical properties of the matrix; and also the optimization of the interactions between the
functional monomer and the template molecule. Hence, a wide range of applications involving
MIPs have been developed, including, solid phase extraction [178], chemical sensors [179],
biomolecules separation [180] and controlled (targeted) drug delivery [181].
Among others, there are two main different ways to perform the imprinting of the complementary
cavities that includes, bulk imprinting and surface imprinting [182]. The "bulk approach"
Chapter 2
34
constitutes the most usual technique employed to imprint small molecule templates. It consists
in a one-step procedure where a pre-polymer mixture is prepared with the target molecule and
assembled on the surface of the transducer. Typically, all the components interact in solution,
resulting in a polymeric matrix holding specific binding sites located in a homogenous
distribution. In contrast, "surface imprinting" enabled the possibility to extend this biomimetic
strategy to larger molecules, such as, proteins, by controlling the location of the imprinted
molecules within the network. Thus, the template molecule is firstly assembled on the solid
substrate surface and afterwards, the synthesis of a thin polymer film is finely tuned in order to
allow the consequent diffusion and removal of this template, by creating the binding sites in
proximity with the material's surface.
Another way to categorize the fabrication of MIPs is based on the covalent and non-covalent
interactions between monomer and template [183]. In the beginning of molecular imprinting, the
use of covalent interactions was a guarantee that the binding sites of the monomer could be
employed in the exact stoichiometric ratio to the template molecule, enabling a higher control of
the imprinted cavities. Later on, non-covalent imprinting become quite attractive due to the
flexibility and simplicity of the method of operation, whereas the interactions occur by hydrogen
bonding, ionic interactions, van der Waals. Although this last strategy requires binding site
monomers to be present in large excess to ensure efficient complexation of the template
molecule, creating random binding sites [184], currently the non-covalent route is the most
widely used approach to fabricate MIPs. Interestingly, despite the development and application
of MIP-based functionalities have been extensively spread due to the wide diversity of materials
available, their "smart" incorporation in biosensing platforms is quite more restricted and a huge
growing has been observed over the last two decades, as seen in Figure 2.19.
Figure 2.19: Graphical representation for the number of publications found in ISI Web of Knowledge
related to A) MIP materials (molecularly-imprinted polymer* or molecular imprinting or MIP*) and B) MIP
materials in biosensors (biosensor* and molecularly-imprinted polymer* or molecular imprinting or MIP*), in
search made April 2018.
Chapter 2
35
Therefore, molecular imprinting technology has become today an important tool to design
nanostructured materials with highly tuneable recognition properties. As mentioned before,
MIPs offer a wide range of applications including drug delivery systems [185], stationary phases
[186], solid phase extraction [187] and also biosensing devices [188]. In parallel, several
characterization techniques have been used during the assembly and application of imprinted
materials [189]. Scanning electron microscopy (SEM), transmission electron microscopy (TEM)
and atomic force microscopy (AFM) have been widely employed for morphological
characterization; X-ray absorption and diffraction and X-ray photoelectron microscopy for
structure analysis; nitrogen adsorption for measuring the specific surface area and pore size of
polymers; and nuclear magnetic resonance (NMR), Fourier-transform infrared (FTIR) and ultra-
violet visible (UV-Vis) spectroscopy for screening the interaction between functional monomers
and template molecules.
Among the different ways of designing an efficient MIP-based biosensor, the choice of the
transduction element is an important aspect to acquire the desirable sensitivity. For instance,
QCM sensors have been widely used as synthetic coatings due to their special features, such
as, low cost fabrication, mild conditions of operation and the ability to track extremely low mass
variations [190]. Furthermore, polymeric-based MIPs with luminescent characteristics hold
unique capability as signal transducers, enabling direct assessment of the binding events, which
results in high selectivity and sensitivity of sensor devices [191]. Another strategy commonly
used to fabricate biomimetic materials is through electrochemical transduction, whereas the
main advantages are good adherence to the sensor surface, quick analysis periods and high
control of the material thickness [192].
The use of MIPs as recognition elements towards the assessment of biological molecules was
constantly reviewed, specifically, the effect of experimental conditions including pH, nature of
the buffer and charges/functional groups have been investigated [187][193]. In recent years, the
high stability and specificity of MIPs have turned these a promising alternative to
immunosensors, bringing out new advantages to 8-OHdG detection. These synthetic materials
are obtained by molecularly-imprinted technology and are linked to longer stability properties
and lower production costs than their natural counterparts. In this context, and as far as one
know, a single approach has been presented in the literature targeting 8-OHdG. It consisted in
creating imprinted sensing films using metal chelating agents (used as monomers) cross-linked
with bisacrylamide [194]. The affinity features of these materials towards 8-OHdG were
measured by QCM gold sensing receptor surfaces, but the detection capabilities were not lower
than 0.01 M. In addition, such materials have been reported in 2008-2009, and since then no
significant developments on molecular-imprinting technology have been addressed in the
literature.
As said previously, the ultimate goal of molecular imprinting is to accomplish biomimetic
materials with equal affinity and specificity as the biological ones. One strategy currently used is
the integration of an imprinting approach with electrochemical sensing technique to enhance
Chapter 2
36
both the sensitivity and selectivity of the sensors [195]. Sometimes, in order to achieve the low
detection limits required, MIPs can be further improved by introducing nano-sized structures,
such as, MWCNT [196], Au-NP [197], graphene [198], among others. In parallel,
electrochemical technology coupled with a molecular-based approach have been successfully
applied for the detection of small molecules, such as, melamine [199], glucose [200], dopamine
[201], creatinine [202] and uric acid [203], which constitutes the ultimate evidence that these
biological metabolites can be sensitively quantified in complex biological matrices.
From a practical perspective, the polymeric fraction of the MIP materials is prepared by radical
polymerization, in the presence of suitable monomers, cross-linkers and initiators. While the
conventional MIP materials use chemical reagents to generate radicals and initiate the
polymerization [204], in recent years photo- [179] or electrically-driven [205] polymerization has
been highlighted. For the purposed of an electrochemical sensing, electrochemical
polymerization is the most straightforward approach to produce ultra-thin polymeric films
deposited directly on the surface of a transducer. Electropolymerization enables to finely tune
the thickness of the imprinted film by controlling the number of cycles and the scan-rate of the
current that is applied in the electrode. Moreover, this in-situ technique allows attaching the
sensor film to electrode surfaces of different shapes and sizes and, at the same time, by
selecting a suitable solvent and supporting electrolyte the morphology of the polymeric matrix
can be tailored. Another advantage of using the electropolymerization technique is the ability to
carry out MIP synthesis and biomolecule immobilization in a one-step procedure and
consequently, enabling a cost-effective approach. Recently, a MIP sensor for dopamine
detection based on a chitosan-graphene mixture has demonstrated how bulk imprinting of small
size molecules can be a simple and successful approach [141]. In addition, carbon nanotubes
and supramolecular cyclodextrins have been chosen to modify electrodes for a simultaneous
determination of DNA bases [148].
A wide range of electroactive monomers can be used for electrochemical polymerization,
resulting in either conducting or non-conducting polymers. Nowadays, the most common
electronically conducting polymers are polyacetylene, polyphenylene, polypyrrole,
poly(aminophenylboronic acid), polythiophene, polyaniline and polyethylenedioxythiophene,
among others [205]. Both conducting and non-conducting MIPs hold important advantages and
limitations that should be taken into consideration during the preparation and operation of these
biomimetic sensing devices. Firstly, during the electropolymerization of a conducting MIP, the
growth of the polymer will proceed indefinitely, while for a non-conducting MIP film the polymer
thickness is self-limited until the growth of the polymer insulates completely the underlying
surface of the conductive electrode. So, for instance, if using a surface molecular imprinting
approach controlling the polymer thickness around the target molecule is possible by carefully
choosing the optimized conditions during the electro-deposition. In addition, the thickness of the
MIP film can also have great implications in the response time and performance of the sensor
[206].
Chapter 2
37
Among the several electroactive functional monomers that allow the formation of non-
conducting polymers, phenol-based MIPs have been widely employed as recognition units in
biosensors. Briefly, phenolic compounds constitute an important class of chemical compounds
consisting of a hydroxyl functional group (–OH) attached to an aromatic hydrocarbon ring
structure. Phenolic compounds represent a large group of biological molecules and, specifically,
biomolecules, such as tyrosine and tyramine are the building blocks of numerous natural
products. The oxidation behaviour of phenol and para-substituted phenolic compounds was
investigated by means of CV, DPV and SWV at a GC electrode in different electrolytes, with
different pH values [207]. Moreover, several mechanisms regarding the oxidation of phenol and
phenol derivatives have been widely described and, herein, a proposed electro-oxidation
pathways for phenol is presented (see Figure 2.20) [208].
Figure 2.20: Adaptation of the proposed mechanism of phenol electro-oxidation [208].
As previously mentioned, during the electrosynthesis of MIPs, the compatibility of the functional
monomer against the target-molecule is a crucial issue that also dictates the performance of the
biosensing device. As far as our knowledge goes, only two studies involving a molecular imprint
approach were used to detect 3-NT biomarker [209][210] and, both of these performed the
synthesis of the polymeric matrix by using conventional free radical polymerization.
Interestingly, in order to achieve the required low detections limits, nanostructured materials
such as bimetallic nanoparticles and carbon dots were incorporated during the assembly of the
biosensor. So, novelty approaches are still on-going in order to facilitate the fabrication and
operation of sensing devices designed to be used in early detection context.
Chapter 2
38
2.5 NANOMATERIALS
Among all requirements to develop a suitable POC biosensor, the ability to detect very
low concentrations of the target molecule constitutes the core piece that limits the analytical
performance of that device. Moreover, the sensitivity of any system can be also directly
correlated between the amount of analyte and the strength of the respective output signal [184].
Over the years, nanotechnology has boosted their way towards the development of important
tools for diagnostic and therapeutic applications. So, while searching for different amplification
strategies, like for instance, the use of labels and electroactive molecules, the incorporation of
nanomaterials during the design of electrochemical platforms have greatly enhanced the
capabilities of the biosensing device [211]. A wide variety of nanoscale materials that includes,
metal nanoparticles, semiconductor nanoparticles, quantum dots, carbon nanosized structures
and magnetic nanoparticles (MNP) have been introduced as electrochemical signal amplifiers
for the assessment of circulating biomarkers (see Figure 2.21).
Figure 2.21: Representation of some nanostructured materials used for diagnostic applications.
The key properties of these nanostructured materials can be finely tuned by mainly controlling
their size and surface features. For instance, for drug delivery applications the diameter of the
particle must be well-defined in order to avoid nanotoxicity effects [212][213][214]. In addition,
nanoparticles hold a unique size-dependent characteristic, particularly regarding magnetic and
optical properties enabling an easy manipulation to obtain good detectable signals. Due to their
high reactive surface area and small particle size, nanomaterials have been applied has an
upgrade in amperometric biosensors causing an enhancement of the current analytical
methodologies [215]. Au-based nanoparticles have been the most employed platform for
Chapter 2
39
immobilization of biological molecules toward biomedical applications [109]. Owing to their easy
preparation methods, good biocompatibility, high chemical stability and catalytic activity, Au-
based nanoparticles have found attractive applications for the assessment of circulating
biomarkers, making it highly compatible with novel POC devices [113]. Moreover, other noble
metals like for instance Pt and silver (Ag) have been widely used for the preparation of
nanoparticles mostly due to their good electrical conductivity, being incorporated in the
development of electrochemical biosensors. Another advantage of these nano-based systems
is their ability to enhance the electron transfer rate between biomolecules and the transducer
platform.
Over the years, manufacturing of this type of nanomaterials has been classified as bottom-up
and top-down approaches, as seen in Figure 2.22. Briefly, the bottom-up approach involves
building up from the atom or molecular constituents to meso-level while the top-down approach
is characterized by reducing the dimension of the original size from the bulk materials [216].
Figure 2.22: Adapted scheme of the different states during nanoparticles fabrication.
As mentioned before, carbon nanostructures have been widely employed not only as
amplification signal strategies, but also as a convenient and straightforward approach to
functionalize different kinds of surfaces [217][218]. However, herein the internal arrangement of
these nanomaterials makes it architecture essential for the application, for which they are
designed (see Figure 2.23). For instance, CNTs have been introduced in biosensor devices as
biorecognition elements in different approaches, such as, a single probe [219] or a support
surface onto the transducer platform [196][220] for amplification purposes.
Figure 2.23: Different nanostructures of carbon A) graphene, B) SWCNTs and C) MWCNTs.
Chapter 2
40
Nanoparticle-based assays can undergo at two different ways: (a) the binding of a label-
nanoparticle to the target biomolecule will produce a measurable signal or (b) the nanoparticle
can be directly employed as the transduction material. Besides the broad spectrum of analytical
applications, another great advantage of these nanotechnology-based assays results from their
ability to automation making these ideal for routinely diagnostic tools. Meanwhile, due to
population variations in the expression of a single biological marker and also the lack of a
known biomarker associated to a specific disease, there has been a great need to develop a
biosensor array for multiplexed relevant biomarkers. Under this scope, various electrochemical
biosensing devices coupled with multiplexing systems have been studied for early and
minimally invasive clinical diagnosis [221]. In addition, different immobilization and
biorecognition approaches can be combined and used simultaneously in order to get array-
based systems, as presented in Figure 2.24.
Figure 2.24: Different immobilization methodologies used during the fabrication of a sensor device: A)
sandwich immunoassay approach; B) MIP-based approach; C) labeled nanoparticle approach and D) ink-
based approach.
In general, nanoparticles can display electronic, magnetic and optical properties that
behave in a different way against their bulk counterpart products. Among all the nanostructured
materials mentioned before, magnetic nanoparticles have been extensively applied in order to
immobilize different types of molecules on the transducer surface. Briefly, iron and oxygen are
chemically combined to produce iron oxides, being the three most common forms of iron oxides
Chapter 2
41
found in nature: magnetite (Fe3O4), maghemite (-Fe2O3), and hematite (α-Fe2O3) [222]. The
preparation of magnetic-based materials, such as, Fe3O4, has received special attention for
their incorporation in bio-related applications, like for instance, magnetic resonance imaging
(MRI) [223], hyperthermia anti-cancer theraphy [224], separation [225], magnetic drug targeting
[226], purification of molecules [227] and also environmental remediation applications [228].
Specifically, functionalized magnetic nanoparticles are being widely applied for sample
preparation procedures due to the following advantages [229]:
(i) high extraction efficiency (due to the high surface area-to-volume ratio);
(ii) fast separation;
(iii) facile preparation and surface modification of the extraction phase;
(iv) high selectivity for the target analytes and suitability for complex matrices;
(v) good reusability;
(vi) excellent dispersibility in aqueous solution and easy to operate;
The high surface to volume ratio in parallel to their biocompatible properties, make magnetic
nanoparticles quite appealing to be used as immobilization support. Moreover, the ability to
induce a simply "switch on-off" magnetic moment in the presence of an external magnetic field
can be a great advantage in terms of solution manipulation, separation, pre-concentration and
even recovery. As shown in Figure 2.25, iron oxide magnetic nanoparticles can be prepared by
three different routes: chemical, physical and biological [222].
Figure 2.25: Adapted graphic concerning the different routes used for the synthesis of iron oxide magnetic
nanoparticles [222].
In general, the most common method to prepare magnetic nanoparticles is by using a chemical
approach, in which iron oxides are synthesized through the co-precipitation of Fe2+
and Fe3+
,
with the addition of a strong base. In contrast with the physical methods that include complex
procedures and show some variability related to the nano-size of the particles, the chemical
method is quite simple and easy-to-use, enabling to finely tune the size, composition and shape
of the magnetic nanoparticle by controlling some experimental parameters, like, type of salt,
Fe2+
and Fe3+
ratio, pH, and ionic strength [230][231][232].
Chapter 2
42
However, magnetic nanoparticles are not completely stable at normal physiological conditions,
showing a great tendency to aggregation due to their hydrophobic nature. Among the different
approaches available to enhance the dispersibility and stability of these iron oxides (see Figure
2.26), silica coating has been one of the most attractive and common technique to be
implemented. With this approach, silica shells not only prevents the nanoparticles from
agglomeration but, in addition, allows to modify and encapsulate other materials enhancing the
versatility and utility of this kind of core-shell systems.
Figure 2.26: Silica applications conjugated with magnetic nanoparticles as nanoplatforms.
Chapter 3
43
CHAPTER 3
3 8-Hydroxy-2'-deoxyguanosine biomarker detection
down to picoMolar level on a plastic antibody film
The results presented in this chapter were published in Gabriela V. Martins, Ana C. Marques,
Elvira Fortunato, M. Goreti F. Sales, "8-hydroxy-2′-deoxyguanosine (8-OHdG) biomarker
detection down to picoMolar level on a plastic antibody film", Biosensors and Bioelectronics
(2016), 86, p. 225-234. doi: 10.1016/j.bios.2016.06.052.
Chapter 3
44
3.1 INTRODUCTION
In the field of biosensors, one of the most important aspects concerning the assembling
of MIPs is the proper integration between transducer and recognition elements. In this sense,
electropolymerization approach enables the formation of selective binding sites, at a precise
spot, closer to the electrode surface. Thus, by controlling some experimental parameters, such
as, the initial concentration of the monomer and also the ratio template-monomer one can finely
tune the performance of the MIP-based sensor. In addition, electrochemistry is a simple, low
cost and quite sensitive tool, easily applied to electroactive species in aqueous media.
Herein, special emphasis has been given to electropolymerization of phenol for the preparation
of MIP-based sensors, due to its straightforward preparation and its ability to interact with
different analytes, through hydrogen bonding and ππ stacking [177][233]. An overview of the
literature have showed that the electropolymerization of various phenolic compounds have been
widely investigated in different electrode surfaces including vitreous carbon electrode [234], Pt
[235], Au [236], carbon steel and stainless steel [237]. The preparation of MIPs of small
molecules have been acknowledged as an easy approach to synthesize biomimetic materials
holding specific functionalities. Combined with an electrochemical approach, MIP-based
sensors have been designed with high sensitivity and selectivity for the detection of small
molecules, like, melamine [238], theophylline [239], creatinine [240].
In the present work, a simple and sensitive electrochemical MIP-based sensor for detection of
urinary 8-OHdG has been assembled via electropolymerization. To this end, 8-OHdG was
employed as the template molecule and phenol as the functional monomer (see Figure 3.1). As
mentioned before, one of the main advantages of the application of polyphenol films is their
facile preparation by electropolymerization in mild aqueous media suitable for biological
molecules. Several experimental parameters have been carefully optimized and the
electrochemical performance of the designed MIP sensor was investigated by CV and EIS.
Moreover, it was employed to detect 8-OHdG in urine samples as a non-invasive approach to
assess the extent of DNA oxidative damage. Thus, the proposed biosensor provides a highly
selective tool to be implemented as an easy-to-use protocol for sensitive detection of 8-OHdG in
biological samples.
Figure 3.1: Schematic representation of the assembly of the gold-modified imprinted sensor.
Chapter 3
45
3.2 EXPERIMENTAL SECTION
3.2.1 Reagents and Materials
All chemicals were of analytical grade and used as supplied without further purification.
All buffer solutions were prepared in phosphate buffered saline (PBS, 0.01 M, pH 7.4) with
ultrapure water Milli-Q laboratory grade. The exact pH values were measured with a pH meter
(Crison Instruments, GLP 21 model). Potassium hexacyanoferrate III (K3[Fe(CN)6]) and
potassium hexacyanoferrate II (K4[Fe(CN)6]) trihydrate were obtained from Riedel-de-Haen; 3-
mercapto-1-hexanol, phenol (for molecular biology) and 8-OHdG (98%) from Sigma-Aldrich;
ethanol absolut (99.8%) from Panreac and the Fluorescein Isothiocyanate (FITC) labeled
antibody against 8-OHdG (polyclonal, 100 g, 0.5 mg/mL in PBS) from Biorbyt. Piranha
solution was prepared by carefully mixing concentrated sulfuric acid (H2SO4, 95%, Normapur)
and hydrogen peroxide (H2O2, 30%, Scharlau) in 5:1 ratio. All experiments were carried out at
ambient temperature.
3.2.2 Apparatus
Electrochemical measurements were performed by using a classical three-electrode
system consisting of a gold-modified electrode as the working electrode, a platinum wire as the
counter electrode and a Ag/AgCl wire as the reference electrode. The diameter of the working
electrode was 2 mm. The electrochemical measurements were conducted with a potentiostat/
galvanostat from Metrohm Autolab and a PGSTAT302N with a FRA module, controlled by
ANOVA software.
FTIR, Raman spectroscopy and SEM characterization was conducted using gold-screen printed
electrodes (Au-SPE) purchased to DropSens (DRP-220AT), instead of the conventional gold
electrode. FTIR measurements were performed using a Thermo Scientic Smart iTR Nicolet
iS10, coupled to the Attenuated Total Reflectance (ATR) smart accessory, also from Thermo
Scientific. For Raman analysis, we have used a Thermo Scientific DXR Raman microscope
system with a 100 mW 532 nm excitation laser. For both FTIR and Raman measurements, data
analysis was performed with OMNIC software. Surface morphology of the polymeric films was
examined in a Carl Zeiss AURIGA Crossbeam SEM-FIB workstation.
The images concerning MIP and NIP labeled with FITC-anti-8-OHdG were collected by a
Confocal Laser Scanning Microscope (LSM 700/ Carl Zeiss). Afterwards, the image analysis
was performed using ZEN 2.1 software (Carl Zeiss).
3.2.3 Gold electrode cleaning
Before use, the bare gold electrode was cleaned by dipping in a mixture of piranha
solution for 20 min, rinsing abundantly with ultra-pure water and drying. Then, the electrode was
carefully polished with aqueous alumina slurries, with successive decrease in particle size
(10.05 m), followed by sonication in ethanol-water mixture. Finally, the electrode was
Chapter 3
46
abundantly washed with ultra-pure water and allowed to dry at ambient temperature. Before
each experiment, the electrode was subjected to cyclic sweeping between 0.2 and +1.5 V, in
0.5 M H2SO4 solution, until a stable cyclic voltammogram was obtained (more or less 5 cycles).
3.2.4 Sensor fabrication
Initially, Au surface of the electrodes was modified with a monolayer of 3-mercapto-1-
hexanol. The clean Au electrode was immersed in a 20 mM thiol solution for 2 hours, at 25 ºC,
in order to form the covalent attachment of thiol to the Au electrode surface. This incubation
step was responsible for the formation of a stable self-assembled monolayer on the electrode
surface through a strong gold-sulfur interaction.
MIPs were prepared by bulk polymerization and all experiments were carried out at ambient
temperature. Previously, the phenol solution was deoxygenated by bubbling nitrogen gas for 15
min. Afterwards, the electropolymerization was performed by CV (3 cycles) in the potential
range +0.1 to +0.9 V, at a scan rate of 20 mVs-1
, in a 0.01 M PBS solution, containing both
phenol monomer and the template molecule 8-OHdG. MIP solutions were renewed every 3
experiments. Then, template removal was carried out by consecutive immersion in ethanol and
PBS solutions, for 30 min each, leading to the formation of recognition cavities in the MIP
structure. This solvent was chosen to ensure an efficient removal of 8-OHdG, while keeping
mild conditions and avoiding pH changes upon the polymeric film.
Control electrodes (NIPs or non-imprinted polymer) were prepared by following the same
procedure but without the presence of the template molecule. The NIP electrode had the same
treatment as the MIP sensor, in order to ensure that variations were only attributed to the
imprinting features. All the modified electrodes were stored at 4 ºC before further use.
3.2.5 Electrochemical assays
Electrochemical measurements for characterization of the modified electrodes were
performed by using different electrochemical techniques, such as, CV and EIS. For CV assays,
the potential was scanned from 0.1 to +0.4 V, at 50 mVs-1
, in 5.0×10-3
mol/L K3[Fe(CN)6] and
K4[Fe(CN)6], in 0.01 M PBS solution, pH 7.4. EIS assays were conducted with the same redox
couple [Fe(CN)6]3-/4-
, at a standard potential of +0.15 V, with a number of frequencies equal to
50, logarithmically distributed over a frequency range 10-3
-104 Hz. All experiments were
conducted in triplicate at ambient temperature.
Calibration curves were made with 8-OHdG standard solutions ranging from 0.10 and 100.0
pg/mL. All standard solutions were freshly prepared in PBS pH 7.4. The time given for 8-OHdG
incubation before reading of the redox probe was set to 20 min.
Chapter 3
47
3.2.6 Surface analysis
Each step of chemical modification on the gold electrode was followed ex-situ by FTIR-
ATR, Raman and SEM analysis. For this purpose, the polymeric films were grown on Au-SPE
for 10 CV cycles. Before measurement, samples were left drying at room temperature, for at
least one day. Regarding FTIR analysis, the infrared spectra were collected after background
correction, with a number of scans set to 500. Raman spectra were recorded using a 4 mW
power, exposure time 60 s, number of exposures 10 and 50 m slit aperture. SEM images were
obtained by using an accelerating voltage of 5 kV with an aperture size of 30 m.
3.2.7 Preparation and characterization of the FITC-labeled surfaces
Fluorescence microscopy was employed to assess the presence and distribution of the
rebound 8-OHdG molecule via fluorescent emission from the FITC labeled antibody. The
fluorescence studies were conducted on MIP and NIP surfaces built on Au-SPE. Afterwards, the
rebinding of 8-OHdG molecule was performed in the same conditions of the calibration
procedure. The antibody was previously diluted 1:100 in PBS and incubated on top of the
electrodes for one hour (4 ºC). All the fluorescence measurements were carried out with the
excitation wavelength of 488 nm and the emission wavelength of 520 nm. Control images of
non-imprinted electrodes were obtained in identical experimental parameters. Afterwards, all
images were processed in equal conditions, by applying a brightness filter of 80% and a
contrast filter of 90% in order to improve fluorescence visualization.
3.2.8 Selectivity studies and analysis in urine samples
The selectivity experiments were carried out through incubation of the MIP-based sensor
in 8-OHdG solution (5 pg/mL) in the presence of each individual interfering species. Uric acid,
citric acid and glucose were chosen as interfering molecules and used at concentrations
mimicking the physiological levels. In addition, selectivity features of the MIP were assessed
directly in human urine samples, due to their complex nature as a biological fluid. The urine
samples were collected in sterile bottles to avoid any contamination. After collection, the fresh
urine samples were directly frozen in aliquots of 1 mL. Before analysis, the samples were
diluted in PBS buffer, in a 1:1000 ratio.
The detection analysis of 8-OHdG in urine samples was tested by immersing the MIP sensor in
the diluted samples during a period of incubation of 20 min, followed by EIS analysis.
Preliminary recovery tests were performed by adding a known concentration of 8-OHdG to urine
samples.
Chapter 3
48
3.3 RESULTS AND DISCUSSION
3.3.1 Optimization of experimental variables
During the fabrication of MIP materials, several parameters need to be carefully
optimized. Herein, the concentration of monomer, the number of CV cycles and its ratio against
the target analyte have been considered, as these are found critical steps at the MIP film
assembly [201][241]. The pH was also a very important parameter, but it was kept at this stage
equal to 7.4, as this is a close condition to that in biological fluids.
The electropolymerization of phenol was achieved by CV. Figure 3.2A illustrates the first
voltammetric cycle concerning the electropolymerization of phenol, at pH 7.4, for different
concentrations of monomer. As shown, pure phenol solutions exhibited only an oxidation peak
around +0.6 V, indicating an irreversible oxidation reaction of the monomer on the electrode
surface. It was clear that the peak current increased with increasing monomer concentration
0.25 < 0.50 < 1.25 mM. Due to the non-conductive behaviour of polyphenol [205], this
accounted the growth of a strongly passivating polymer layer on the electrode surface. In
general, a great increase in the overall resistance of the sensing layer may hinder the sensitivity
of the device. On the other hand, for concentrations of monomer below 0.25 mM, the efficiency
of the polymerization reaction may be small, questioning the stability of the polymeric film.
Therefore, 0.25 mM was identified as the optimum phenol concentration for the preparation of
the 8-OHdG sensor.
Figure 3.2: A) Cyclic voltammograms of a gold-modified electrode immersed in 0.01 M PBS aqueous
solution containing different concentrations of monomer phenol (0.25, 0.5 and 1.25 mM), pH 7.4, scan rate
20 mVs-1; B) Charge variation during electropolymerization of phenol (3 cycles) obtained from MIPs with
different ratios of template to monomer (1:3 and 1:1) and NIP in 0.01 M PBS.
Another challenge within the MIP technology is to find an agreement regarding the
concentration and distribution of recognition sites close to the sensor surface and,
simultaneously, well connected along it [242]. Likewise, the number of cycles during the
Chapter 3
49
electropolymerization of the monomer can tailor both the thickness and structural characteristics
of the resulting polymeric film [236]. In our investigations, no more than 3 voltammetric cycles
were used during the electropolymerization of phenol to avoid that 8-OHdG molecules could be
buried deep within the polymeric film, inhibiting their subsequent elution in order to create
effective recognition sites. In addition, a higher number of cycles would generate a less
conductive surface and, subsequently, a less sensitive device.
Next, the effect of the mole ratio of 8-OHdG molecule to phenol monomer was also
investigated. Figure 3.2B presents the time-dependent charge response during
electropolymerization of NIP films and MIP films with different ratios of template/monomer (1:3
and 1:1). The thickness of the polymer film can be roughly estimated from the total charge
during electropolymerization [243]. According to Figure 3.2B, the total charge resulting from
phenol electro-oxidation in the NIP is around 210-4
C and, therefore, the thickness of the
polymeric film was about ~60 nm. This result is in agreement with previous studies performed in
similar conditions [236]. Different ratios of template to monomer were investigated for the MIP
material, in order to optimize the amount of binding sites available for the selective re-binding of
8-OHdG. When the ratio was 1:3, the charge response decreased comparatively with NIP, as
shown in Figure 2B, while in the case of a ratio 1:1, the charge response increased again. So,
although our results have showed that the presence of the target analyte affects the total charge
resulting from phenol oxidation, the previous approach used to estimate the thickness of the
films can only be applied to the polymer layer. Overall, the presence of the target molecule
within the film hindered the growing of the polymer, as it did not participate in the polymerization
stage. In addition, calibration curves of 8-OHdG in PBS solution at pH of 7.4 were performed for
both MIPs and the highest electrochemical response (highest sensitivity) was obtained for the
MIP with the ratio template to monomer of 1:1(see Figure 3.3). Thus, the optimal template-
monomer ratio of 1:1 was chosen for the following studies expressing a final concentration of
imprinted molecule of 75 g/mL.
Figure 3.3: Calibration curves of 8-OHdG obtained for MIPs with different ratios of template/ monomer, 1:3
(closed gray circles) and 1:1 (open black circles).
Chapter 3
50
3.3.2 Preparation and electrical follow-up of MIP sensor
Figure 3.4: Cyclic voltammograms concerning the electropolymerization of 0.25 mM phenol in 0.01 M
PBS, pH 7.4, (scan rate 20 mVs-1, 3 cycles) at gold-modified electrodes with (dashed line) and without
(straight line) the template molecule 8-OHdG.
Figure 3.4 shows a typical voltammogram recorded during the electropolymerization of
phenol in the presence (dashed line) and absence (solid line) of 8-OHdG. The only peak visible
occurred around +0.6 V and is due to phenol oxidation. This record also indicated that 8-OHdG
was not electroactive in the positive range 0.10.9 V, meaning that the template structure was
not electrochemically altered and was preserved during the imprinting step. Furthermore, it was
clear that the current decreased with increasing number of cycles and the highest current was
obtained in the 1st cycle. As expected, after the 2
nd and 3
rd cycles, the phenol oxidation peak
started to disappear, which resulted from the formation of a non-conductive layer that blocked
the access of the monomers to the electrode surface. It was interesting to see that the peak
current in the presence of 8-OHdG was smaller than that obtained without it, which can be a
strong indication of the existing interactions between the phenol monomer and the template
molecule. The hydroxyl groups in the 8-OHdG molecule could be interacting with the phenol
monomer through hydrogen bonding. This is consistent with several other reports that have
confirmed the importance of electrostatic interactions and hydrogen bonding between template
molecules and phenol [53][197]. Recently, the construction of graphene-based MIPs have
supported that hydrogen bonds between template and polymeric matrix can enhance the
recognition and selectivity for the target molecule [193]. Moreover, previous studies have
demonstrated that π donor-acceptor interactions between electropolymerized phenol and an
imprinted template molecule can be used as a novel method to generate imprinted polymers
[233].
Chapter 3
51
Figure 3.5: A) CV of the gold electrode (green line), thiol-modified gold electrode (red line), NIP and MIP
after electropolymerization (blue and grey line, respectively) and after template removal (black lines, on the
right side), measured in aqueous solution containing 5 mM [Fe(CN)6]3-/4-
in 0.01 M PBS pH 7.4 and B) EIS
of (a) gold electrode, (b) thiol-modified gold electrode, NIP (c) before and (d) after removal, MIP (e) before
and (f) after removal, in aqueous solution containing 5 mM [Fe(CN)6]3-/4-
in 0.01 M PBS.
The construction of NIP and MIP sensors was followed in-situ by CV and EIS measurements in
5 mM [Fe(CN)6]3-/4-
solution, prepared in 0.01 M PBS as the supporting electrolyte. Figure 3.5A
shows the cyclic voltamogramms obtained with gold-modified electrodes, where a couple of
well-defined redox peaks was evidenced, as a result of the reversible electron transfer of the
redox pair [Fe(CN)6]3-/4-
. After pre-incubation of thiol onto the gold surface, the current of the
redox peak observed decreased, meaning that this modification slightly increased the electrical
resistance of the working electrode. This outcome was expected as the result of the
spontaneous formation of a closely packed monolayer via a strong gold-sulphur interaction
between the SH group and the gold. Afterwards, it was found that the formation of polyphenol
polymeric layer hindered the electron transfer process resulting in a good blocking efficiency on
the electrode surface. Thus, the NIP and MIP formation are characterized by the disappearance
of the pair of redox peaks [244]. Interestingly, the MIP film showed lower current change
throughout the voltammogram in comparison to the NIP, which was a strong evidence of the
Chapter 3
52
presence of the template 8-OHdG within the polymer matrix. Once the template was extracted
from the MIP, the film recovered some of the current change lost, accounting the exit of the
imprinted template from the polymer matrix and thereby confirming the formation of binding
sites. As expected, after the 8-OHdG removal step (with ethanol-PBS solutions), the
electrochemical behaviour of the NIP was unaltered.
EIS studies were used to follow-up the variation of gold-modified electrodes after each chemical
change, since this electrochemical measurement holds a high sensitivity and it is strongly
suitable in the detection/follow-up of small alterations of non-conductive polymers. Randle's
equivalent circuit was adopted to model the physiochemical process occurring at the gold
electrode surface. In general, the impedance spectra included a semi-circular portion at higher
frequencies and a linear portion at lower frequencies which corresponded to the electron-
transfer resistance and the diffusion process, respectively. The diameter of this semicircle is the
Rct that controls the electron transfer kinetics of the redox-probe at the electrode interface.
Figure 3.5B illustrates the Nyquist diagrams of the electrodes fabricated at each step in the
presence of 5 mM [Fe(CN)6]3-
/[Fe(CN)6]4-
, which were carried out at the formal potential of 0.15
V with a frequency range of 10000 to 0.001 Hz. The bare gold surface presented a small
semicircle domain as the result of a very fast electron-transfer process (Figure 3.5B-a).
Afterwards, the modification with the thiol monolayer onto the gold surface gave rise to a
subsequent increase in Rct (Figure 3.5B-b). As expected, the electropolymerization of phenol
resulted in a substantial increase of impedance (Figure 3.5B-c) due to the blocking of electron
transfer by the polymeric matrix. Furthermore, a higher increase of the impedance was
observed when the polymerization of phenol took place in the presence of 8-OHdG (Figure
3.5B-e), accounting the presence of 8-OHdG entrapped inside the polymeric matrix (consistent
with the fact that the addition of 8-OHdG promotes an Rct increase, evidenced in Figure 3.3.
Finally, the elution of the template from the polymer matrix is acknowledged as one of the most
important steps, since it holds a direct influence on the sensitivity to the template recognition.
Previous studies with other small template molecules, like 8-OHdG, have performed the
removal of these molecules from the MIP structure by extraction with PBS or ethanol solutions
[201][241]. For instance, a capacitive sensor developed by electropolymerization of phenol has
achieved theophyline template elution via aqueous-ethanol mixtures [244]. Herein, the removal
of the template was achieved by incubating the MIP film in ethanol and PBS. This procedure
has led to a substantial Rct decrease of the MIP film (Figure 3.5B-f), accounting the formation of
imprinted cavities that facilitate the diffusion of Fe(CN)63-/4-
through the polymer network. The
same treatment applied to the NIP film has generated only a slight Rct reduction, correlated to
the washout of unreacted monomer or small oligomers (Figure 3.5B-d). Thus, it seems evident
that the huge Rct decrease at the MIP layer is mostly linked to the successful exit of the target
template from the polymeric material.
Chapter 3
53
3.3.3 Characterization of the modified surfaces
FTIR and Raman measurements were conducted in order to verify each chemical
modification step along the sensor assembly, specifically, the formation of polymeric films
produced by electro-oxidation of 0.25 mM phenol in PBS solution. In both assays, the analysis
was performed directly on Au-SPE. Figure 3.6 shows the FTIR-ATR and Raman spectra of bare
Au, the thiol modification, and NIP and MIP films assembled on the working electrode area.
Figure 3.6: A) FTIR-ATR spectra of gold, thiol-modified gold, NIP and MIP electrodes; B) Raman spectra
of gold, thiol-modified gold, NIP and MIP electrodes and C) typical image from Raman, measured at 50x
magnification, of the gold-screen printed electrodes (Au-SPE).
Regarding the FTIR data, the assembling of a thiol layer on the gold surface was confirmed by
the appearance of a well define band around 1100 cm-1
that could be attributed to the stretching
vibration of the CO bound [235]. The formation of the polymeric layer of polyphenol was
confirmed by the broad band associated with the aromatic out-of-plane CH deformation
vibrations in the range 750-650 cm-1
(only visible in NIP and MIP spectra) [245]. No differences
could be verified on the FTIR spectra of NIP and MIP materials, but this was consistent with the
fact that their main chemical composition was the same.
The Raman spectra of the Au-SPE showed 3 bands at 500, 550 and 830 cm-1
that could be
assigned to carbon compounds that are present in the manufacturing of the gold ink paste.
Chapter 3
54
Comparison between Au-SPE and thiol-modified spectra showed a substantial reduction in the
intensity of the 830 cm-1
peak as the result of the assembling of a monolayer of thiol, thus
covering the gold surface. Spectra of NIP and MIP films produced over the Au-modified-SPEs
presented a significant increase of the Raman shift at 500 and 550 cm-1
, that could be attributed
to the aromatic ring deformation, confirming the growth of the polymeric chain. Interestingly, the
two broad undefined bands around 1500 and 2900 cm-1
are probably a strong indication of
surface modification due to polymerization, as these are expected to be related with the
emission of fluorescence coming from the aromatic rings.
Figure 3.7: A) SEM micrographs of NIP and MIP electrodes and B) confocal imaging of FITC antibody
against 8-OHdG attached to NIP and MIP surfaces.
Surface morphology of the prepared NIP and MIP electrodes was further studied by means of
SEM. From figure 3.7A, we verified that the formation of the phenol polymer seems to be well
distributed/spread on the surface. As expected, due to the small dimension of 8-OHdG
molecule, it was impossible to visualize the imprinted cavities but, interestingly, MIP surface
seems to hold a more globular morphology, with more islets, compared with the NIP one.
Moreover, the comparison between NIP and MIP images have showed that the MIP surface
presents a higher porosity than NIP. This different behaviour in the polymer morphology can be
another strong indication that the presence of 8-OHdG molecule during electropolymerization
resulted in a different structural polymeric growth.
Previous studies have reported the design and development of aqueous phase molecular
imprinting coupled to confocal microscopy imaging, aiming to visualize the imprinting effect
[246]. Herein, we have followed the binding of a FITC labelled antibody targeted against 8-
OHdG to the surface of the sensor in order to track the 8-OHdG oxidative stress biomarker
rebound into the imprinted cavities. Figure 3.7B presents confocal micrograph images of FITC-
labelled antibody 8-OHdG on the electrode surfaces. The well-localized fluorescence signals
Chapter 3
55
present on MIP sample indicated that the FITC labelled antibody have selectively bound to the
target molecule. Moreover, fluorescence signal points were localized in a homogeneous manner
only on the MIP electrode, which is strong evidence that the target molecule was rebound and,
consequently, imprinted cavities have been successfully created. In parallel, the image of a non-
imprinted polymer control taken at identical experimental parameters presents a distinct lack of
a homogeneous distribution of fluorescence. Furthermore, our findings also confirmed that non-
specific labelling almost did not occur. This approach has already been acknowledged as a
useful tool to assess the specific and selective attachment of target molecules for sensoring
purposes [247].
Overall, the combined information of FTIR and Raman spectra, the electron microscopy images
and confocal microscopy photographs confirmed the successful formation of NIP and MIP films
and the ability of MIP film to rebind selectively and sensitively to its target compound.
3.3.4 Performance of MIP sensor
3.3.4.1 Calibration curve
The analytical performance of 8-OHdG sensory materials was evaluated by recording
calibration curves. Figures 3.8A-B presents EIS calibration curves plotted with the Rct of MIP
and NIP sensors against the logarithm concentration of 8-OHdG. For each calibration study,
sensors were previously incubated in PBS solution until a stable response was obtained.
Afterwards, the time given for 8-OHdG incubation was set to 20 min.
In order to evaluate the specificity and selectivity of the assembled biosensors, the value of the
imprinting factor (IF = [template rebound by MIP] / [template rebound by NIP]) can be used as a
comparative tool. From the ratio of sensitivity of the MIP sensor to 8-OHdG and to the NIP one,
the IF was determined to be ~6.3.
Figure 3.8: A) Nyquist plot of MIP sensor in 5 mM [Fe(CN)6]3-/4- in 0.01 M PBS pH 7.4, previously
incubated in increasing concentrations of 8-OHdG and B) the corresponding calibration curves for both
MIP and NIP sensors; C) Calibration curves of NIP and MIP sensors for different 8-OHdG concentrations
in urine samples, measured in 5 mM [Fe(CN)6]3-/4- in 0.01 M PBS pH 7.4. All error bars represent the
standard deviation for three independent measurements.
Chapter 3
56
The results obtained demonstrated that the resistance of the sensing layer increased after
incubating an 8-OHdG solution in the MIP film, which could be due to the re-binding of 8-OHdG
onto the imprinted cavities that hindered the electrical features of the sensing surface
[Fe(CN)6]3-/4-
. In general, increasing concentrations of 8-OHdG increased the diameter of the
semicircles in the Nyquist plot (Rct), indicating that 8-OHdG bound to the sensory layer
increased the charge-transfer resistance of the probe. The calibration curve obtained for the
MIP sensor showed a linear relationship over 8-OHdG concentration in the range [0.1100]
pg/mL. The limit of detection (LOD), was 0.74 pg/mL, calculated by extracting the first standard
from the calibration curve and extending the linear ranges observed, as in potentiometric
devices that respond to concentration in a logarithm basis. The response of the control
electrode (NIP) was independent of the 8-OHdG concentration and it was kept at very low
values for all concentrations within that range (random and small Rct values). All the results
were normalized against the blank value (PBS). Furthermore, the reproducibility of the sensors
for quantification of 8-OHdG was investigated over the entire linear range and the results
showed that the relative standard deviation (RSD) was 0.5-8.8 %.
3.3.4.2 Selectivity studies
Uric acid is typically identified as the major electrochemical interfering species in
biological samples, mainly due to its relatively high concentrations, and, here specifically, the
similar structural characteristics to the 8-OHdG molecule. Therefore, in order to verify the
selectivity of the sensory device, uric acid, citric acid and glucose were selected as interfering
species. Data presented on Figure 3.9 showed that these molecules almost have no
interference in the determination of 8-OHdG, showing variations of 3.0, 12.0 and 0.5% for uric
acid, citric acid and glucose, respectively. Moreover, to assess its effect upon the biosensor
response, the calibration curves were performed directly in human urine samples (uric acid
levels of ~500 g/ml). EIS calibration curves against 8-OHdG concentration for MIP and NIP
sensors are presented in Figure 3.7C. It was interesting to observe that the MIP maintained the
linear EIS response over the considered concentration range, with just a small decrease of the
sensitivity. In contrast, the NIP sensor presented a quite random behaviour which can be a
strong indication that its response is caused by non-specific adsorption at the electrode
modified-surface. Thus, the proposed sensor showed good analytical performance in terms of
sensitivity, selectivity and rapid response towards 8-OHdG determination.
Chapter 3
57
Figure 3.9: EIS measurement of MIP-based sensor recorded after incubation in 5 pg/mL 8-OHdG solution,
alone and in the presence of uric acid (0.4 g/mL), citric acid (0.5 g/mL) and glucose (0.1 mg/mL). All
solutions were prepared freshly on PBS pH 7.4.
3.3.4.3 Analysis of spiked human urine samples
Urine samples were assayed and recovery experiments were carried out via the standard
addition method by adding 0.25, 2.5 and 50 pg/mL. Assuming a null concentration of the urine
sample (because it was used as background media for the overall calibration), the obtained
recovery values were of 92.6, 111.6 and 110.3%, respectively. But, in biological samples 8-
OHdG molecule is present in vestigial levels and so, herein, we also estimated this "background
concentration". Thus, we have employed the Gran's method of multiple standard addition to
estimate the original concentration of the 8-OHdG in urine samples [248].
Figure 3.10: Calibration curve of 8-OHdG in a urine sample. Rct relative corresponds to the normalized
value of charge transfer resistance against the PBS measurement for each spiked level and S is the slope
of the experimental calibration, obtained from three independent measurements.
Chapter 3
58
As can be seen in Figure 3.10, the plot of 10(Rct relative/Slope)
versus the concentration of 8-OHdG
(added) results in a linear behaviour where the x axis interception is a direct indication of the
unknown 8-OHdG concentration initially present in the urine real sample (before spiking). The
calculated 8-OHdG concentration in the real urine sample (before 1:1000 dilution) was 4.1
ng/mL, a value that still is below the maximum limit in healthy humans.
Overall, these results demonstrated that the method was suitable for the determination of the
total content of 8-OHdG in urine samples. Although there are already few researches on the
application of sensitive sensors for detection of 8-OHdG, these often require the sample pre-
treating stages to eliminate the presence of interfering species [249].
3.4 CONCLUSIONS
In the present study, an electropolymerization method was selected for the synthesis of a
MIP-based electrochemical sensor targeted for 8-OHdG recognition and detection. Under this
approach, phenol was electrochemically deposited on gold-modified electrodes in the presence
of the template molecule 8-OHdG. The control of some experimental parameters of the
electropolymerization reaction enabled the preparation of thin homogenous films, which
constitute an important issue in order to achieve trustable, accurate and repeatable sensor
responses. Our results demonstrated that 8-OHdG molecule was successfully entrapped into
the polymeric matrix, enabling a three-dimensional structure with numerous imprinted cavities
sites. The developed electrochemical biosensor showed high sensitivity and selectivity towards
8-OHdG over the wide concentration range. The successful application of the proposed sensor
in the analysis of human urine samples has evidenced its promising features in the early
diagnosis of cancer in point-of-care.
Overall, a simple and easy-to-go detection method of 8-OHdG in biological samples was
possible without any previous sample preparation, providing a great advantage for this new
analytical methodology. Compared to previous methods (see Table 3.1), the work described
herein showed the best lower limit of linear range and limit of detection, coupled with a wide
range of concentrations of linear response.
Chapter 3
59
Table 3.1: Comparison of the main characteristics of some reported assays used in the detection of 8-
OHdG.
Electrodes Detection
method
Linear range
(nM)
LOD
(nM) References
Dendrimer/Aua HPLC-ECD
b __ 1.2 [26]
Nafion/SWCNTsc/GCE
d DPV
e 30-1250 8 [27]
P3MTf/GCE CV
g 700-70000 100 [250]
ECLh immunosensor/
CQDi/Au/SiO2
ECL 0.7-700 0.3 [146]
GOj/Nafion/GCE LSV
k 70-33040 1120 [251]
PICAl/CHI
m/GCE
immunosensor DPV 0.35-35000 0.11 [252]
Metal-Chelate/
Au/QCMn
QCM 10.0-3500 7.5 [194]
ssDNAo/GNs
p/GCE CV 5.6-36155 0.875 [249]
Immunosensor/Au EISq/ SWV
r 0.071-25 __ [253]
EPPGEs SWV 500-100000 28 [124]
Pht-MIP/Au EIS 0.0035-3.5 0.0074 This work
aAU: Gold;
bHPLC-ECD: High Performance Liquid Performance with Electrochemical Detection;
cSWCNTs: Single-Walled Carbon Nanotubes;
dGCE: Glassy Carbon Electrode;
eDPV: Differential Pulse
Voltammetry; fP3MT: Poly(3-methylthiophene);
gCV: Cyclic Voltammetry;
hECL:
Electrochemiluminescence; iCQDs: Carbon Quantum Dot
;jGO: Graphite Oxide;
kLSV: Linear Sweep
Voltammetry; lPICA: Poly(indole-5-carboxylic acid);
mCHI: Chitosan;
nQCM: Quartz Crystal Microbalance;
oss-DNA: Single-Stranded DNA;
pGNs: Graphene Nanosheets;
qEIS: Electrochemical Impedance
Spectroscopy; rSWV: Square Wave Voltammetry;
sEPPGE: Pyrolytic Graphite Electrode;
tPh: Phenol.
Chapter 3
60
Chapter 4
61
CHAPTER 4
4 Paper-based sensing device for electrochemical
detection of oxidative stress biomarker 8-hydroxy-2'-
deoxyguanosine in point-of-care
The results presented in this chapter were published in Gabriela V. Martins, Ana P. M. Tavares,
Elvira Fortunato, M. Goreti F. Sales, "Paper-Based Sensing Device for Electrochemical
Detection of Oxidative Stress Biomarker 8-Hydroxy-2′-deoxyguanosine (8-OHdG) in Point-of-
Care", Scientific Reports (2017), 7, p. 14878-14887. doi: 10.1038/s41598-017-14878-9.
Chapter 4
62
4.1 INTRODUCTION
As the most abundant oxidative product of DNA, 8-OHdG detection allows a premature
assessment of cancer disease, thereby improving diagnosis, prognostics and survival rates.
Therefore, the relevance of 8-OHdG as OS biomarker is also confirmed by the numerous
methods published in the literature aiming its determination. A sensitive automated flow
immunosensor has been developed for detection of urinary 8-OHdG at concentrations of 0.05
ng/mL [249]. Also, the combination of sensitive semi-conducting silicon nanowire with the
specificity of immunoassays have resulted in a biosensor device to detect few nanomolar
concentrations of 8-OHdG [254].
Overall, electrochemical sensors are increasingly gaining attention due to their high sensitivity
(low detection limits), small dimensions, low cost, easy automation and operation. Special
importance has been given to carbon electrodes and the use of nanostructured materials, such
as graphene, nanoparticles and carbon nanotubes. These nanomaterials are currently used as
a surface modification approach to accelerate electron transfer and enhance the
electrochemical activity of biomolecules due to their intrinsic characteristics such as, higher
surface area, good conductivity and signal stability [250][251]. Even after such amazing
developments, an urging need for OS assessment in POC remains. Specifically, there is a gap
regarding rapid, portable, inexpensive and simple screening platforms for biomarker analysis.
In turn, paper-based sensors have become a promising platform for lab-on-a-chip devices,
offering high selectivity and sensitivity for application areas such as, health diagnostic, food
quality control and environmental monitoring [131]. In particular, a recent overview about
diagnostic paper-based biosensors has discussed the integration of nanomaterials for the
detection of nucleic acids, proteins and cells [254]. Although there are already very few studies
performed on paper electrode devices for quantitative analysis of 8-OHdG biomarker, these still
require the immobilization of antibodies [255].
Thus, this work aims the development of low cost easy-to-use label-free paper-based,
electrochemical sensor for the determination of OS biomarkers, which targets 8-OHdG as proof-
of-concept. For this, we have investigated the redox behaviour of 8-OHdG on different carbon-
modified surfaces and the optimization conditions for its selective detection in biological
samples. The redox reaction of 8-OHdG is presented elsewhere [147] and the detection
process of the modified electrodes can be found in Figure 4.1. Here, the electrochemical
performance of the biomarker 8-OHdG at the modified-paper electrode was followed by means
of DPV and the several electrochemical and chemical variables optimized and evaluated.
Chapter 4
63
Figure 4.1: Schematic representation of the oxidation process of 8-OHdG molecule followed on a
conductive carbon paper substrate: 1) hydrophobic white paper as substrate; 2) conductive carbon-coated
paper; 3) in-situ electrochemical measurement.
4.2 EXPERIMENTAL SECTION
4.2.1 Reagents and Materials
All reagents were of analytical grade and used without further purification. PBS (0.01 M,
pH 7.4), TRIS (hydroxymethyl)aminomethane (0.01 M, pH 9.1 and 9.5) and acetate buffer (1
mM, pH 5.1) were used as buffer solutions and prepared with ultrapure water Milli-Q laboratory
grade. The pH values were measured with a pH meter (Crison Instruments, GLP 21 model).
Graphite powder (fine extra pure, particle size < 50 m) was purchased from Merck and used as
received. Poly(3,4-ethylenedioxythiophene) (PEDOT) nanoparticles dispersion in H2O, 8-OHdG
(98%), uric acid (>99% crystalline), sulphuric acid (H2SO4, 95-97 %), MWCNT and multi-walled
carbon nanotube, carboxylic acid functionalized (MWCNT-COOH) were obtained from Sigma
Aldrich; poly(vinyl chloride) carboxylated (PVC-COOH) from Fluka; N,N-dimethylformamide
(DMF) from Analar Normapur and ascorbic acid from Riedel-de-Haen. All measurements were
carried out at ambient temperature.
4.2.2 Apparatus
Electrochemical measurements were performed by using a three-electrode system
composed by a carbon-ink coated paper with an electrode area of 20 mm2 as the working
electrode, a Pt wire as the counter electrode and an Ag/AgCl (KCl 3.0 M) wire as the reference
electrode. All the electrochemical measurements including CV and DPV experiments were
conducted with a potentiostat/galvanostat from Metrohm Autolab and a PGSTAT302N with a
FRA module, controlled by ANOVA software.
Chapter 4
64
4.2.3 Fabrication and characterization of the paper-based sensor
The working electrode was constituted by small parts of cellulose paper (5 x 2 mm) hand-
coated with a conductive carbon-based ink. The protocol for the preparation of this conductive
carbon-based surface is described elsewhere [255]. Briefly, graphite powder was doped with
PVC-COOH and dispersed in DMF with magnetic stirring at room temperature. In addition, a
small amount of different nanostructured materials were added (separately) to the previous
mixture and left stirring for several hours before using. Specifically, we have tested PEDOT
nanoparticles dispersed in water, MWCNT and MWCNT-COOH powders. The sensor was
prepared by drop-coating the surface of the hydrophobic paper with the conductive carbon-
based ink. Afterwards, the electrode area was precisely delimited by using paraffin wax.
Carbon-coated paper was characterized by means of Raman spectroscopy. The Raman
spectral analysis was carried out using a Thermo Scientific DXR Raman microscope system
with a 100 mW 532 nm excitation laser (operational conditions: 20 min of photobleach and 5
min of collect time). Data analysis was performed with OMNIC software.
4.2.4 Electrochemical assays
Initially, the working electrode was electrochemically cleaned by performing voltammetric
sweeps between -0.2 V and +1.5 V in PBS pH 7.4, until a stable voltammogram was obtained
(more or less 50 cycles). Before use, sensors were dried and stored at room temperature.
The electrochemical response of 8-OHdG in PBS solution at pH 7.4 was investigated by
performing CV measurements over the potential range of +0.1 - +0.8 V, in order to find the
oxidation potential. During the electrochemical measurements the three-electrode system was
always submersed into 1 mL of sample volume. Afterwards, differential pulse voltammograms
were recorded after pre-conditioning the working electrode in 8-OHdG solution at a specific
potential. The DPV experimental conditions used were a potential range between +0.2 V and
+0.7 V, pulse amplitude of 25 mV, pulse width of 50 ms, scan-rate of 100 mVs-1
and an
equilibration time of 5 s. In order to regenerate the sensors, a cleaning protocol was performed
to remove the analyte adsorbents by immersion of the used sensors in 0.01 M PBS pH 7.4,
followed by (5) successive CV scannings over the potential range 0 - +0.7 V at a scan-rate of 50
mVs-1
.
Calibration curves were performed by DPV analysis for 8-OHdG in the range 50-1000 ng/mL in
PBS solution at pH 7.4 and all the logarithmic scales were calculated from ng/mL values.
Selectivity studies were carried out by competitive assays between 8-OHdG (0.1 mM) and each
interfering specie, with similar concentration. Here, uric acid and ascorbic acid were selected as
interfering molecules due to the fact that they may co-exist with 8-OHdG in biological fluids,
holding also electro-active properties. Additionally, different buffer solutions were prepared at
different pH environments, such as, PBS at neutral pH (pH 7.4, 10 mM), Tris buffer at basic pH
(pH 9.1, 2 mM) and Acetate buffer at acidic pH (pH 5.1, 1 mM). The performance of the sensor
Chapter 4
65
was tested directly in Foetal Bovine Serum doped with 8-OHdG in the concentration range 20-
1000 ng/mL. The serum was previously 1:10 diluted in PBS buffer.
4.3 RESULTS AND DISCUSSION
4.3.1 Electrochemical behaviour of 8-OHdG
Figure 4.2 shows the cyclic voltammograms of 8-OHdG (50 g/mL) on PBS pH 7.4
performed for different scan rates onto the surface of the paper-modified electrodes. Here, a
well-defined oxidation peak appeared around +0.42 V and a small reduction peak was visible at
+0.40 V. Thus, the peak shape of the CV demonstrated a typical quasi-reversible
electrochemical reaction. In agreement with previous electrochemical studies, it was also
observed that the oxidation peak potential shifted gradually towards more positive values with
the increase of scan-rate [153].
Figure 4.2: Successive cyclic voltammograms performed in PBS at pH 7.4 with 8-OHdG molecule at
different scan rates. Inset: calibration plot of the 8-OHdG oxidation peak current versus scan rate.
Additionally, the influence of scan-rates on the electrochemical oxidation of 8-OHdG on these
graphite-coated electrodes was also investigated. As can be seen on the inset figure, a linear
relationship (R2 = 0.994) between the anodic peak current and the scan rate is evident,
suggesting that the electrochemical oxidation process of 8-OHdG is mainly adsorption-
controlled. Our results are in good agreement with other studies performed on various carbon-
modified surfaces [249][251]. Herein, CV was employed to determine the oxidation potential of
8-OHdG to be used in further electrochemical experiments (+0.41 V), a value that is in
agreement with other similar electrochemical sensors [256][257][258].
Chapter 4
66
4.3.2 DPV analysis of 8-OHdG on paper-modified electrodes
Graphitic materials are known for their good structural, electronic, mechanical, optical,
thermal and chemical properties [217]. In the last years, carbon-based nanomaterials, such as,
graphene [259], nanotubes [260], nanoparticles [261] and conductive polymers [262] are being
widely employed for biosensing devices. Due to their high surface area and high electrical
conductivity, these electrochemical sensors have become a potential alternative to the
conventional labour intensive and time consuming assays that are currently used for biomarker
analysis. Since the electrochemical performance of the sensor is mainly affected by coating
characteristics, developing a suitable ink for a specific target should be an important issue to
design novel sensing devices.
Figure 4.3: DPV detection of 200 ng/mL 8-OHdG solution in PBS pH 7.4 on different graphite-based
electrodes prepared after the incorporation of various nanomaterials dispersed in the graphite ink, such as,
PEDOT nanoparticles, CNTMW and CNTMW-COOH.
Under the scope of this work, DPV was chosen as a highly sensitive and selective
electrochemical tool for quantitative analysis, with great application for nucleic acid sensing
[109][151]. Meanwhile, the ability of 8-OHdG to undertake an electrochemical oxidation through
2-electron transfer reaction on carbon surfaces has already been reported [32]. So, in order to
enhance the DPV signal of the oxidation of 8-OHdG on the paper-modified electrodes, we have
incorporated and tested the effect of some highly conductive materials into the graphite-ink.
Figure 4.3 displays the corresponding peaks of oxidation of 8-OHdG on different graphite-based
electrodes, prepared through the incorporation of some conductive materials such as, PEDOT
nanoparticles, CNTMW and CNTMW-COOH. Our data showed that the presence of the
conducting polymer PEDOT seems to greatly improve the electro-catalytic properties of the
substrate leading to an enhancement of the electrochemical response of 8-OHdG in comparison
with the electrode coated with the graphite alone. This outcome can be attributed to the
symbiotic combination of the good electronic mobility of the PEDOT polymer with the high
Chapter 4
67
surface area of the nanoparticles resulting in a facilitated charge transfer for 8-OHdG redox
biomarker. Although nano-based materials are often employed in order to accelerate electron
transference, here the introduction of nanotubes during the ink production did not reflect on a
improvement of the peak current intensity. Hence, the obtained oxidation potentials were quite
similar between the different carbon-based surfaces and so, the graphite ink doped with PEDOT
was chosen for the further experiments.
4.3.3 Characterization of the paper-modified electrodes
Raman spectroscopy has become a popular, powerful and non-invasive tool to
characterize the structural organization of carbon and related materials. Figure 4.4 shows the
Raman spectra for the different graphite-based electrodes prepared with the inks doped with
various nanomaterials.
Figure 4.4: Raman spectra of the different graphite-based electrodes prepared after the incorporation of
nanomaterials dispersed in the graphite ink, such as, A1) PEDOT nanoparticles, A2) Graphite, A3)
CNTMW and A4) COOH-CNTMW, with the calculated ID/IG ratios and B) RAMAN spectra with the
magnification of the D (Disorder) band, in full-scale mode.
Firstly, in all samples, the typical G and D bands appeared at 1581 cm-1
and 1350 cm-1
,
respectively, which is in accordance with previous Raman spectrum for graphite samples
[255][263][264]. The G band is often associated to the stretching of the C-C in graphitic
materials, common to all sp2 carbon systems, and the D band is assigned to the presence of
disorder in the sp2-hybridized carbon system [265]. In addition, another peak was also clearly
visible around 2700 cm-1
that is assigned to the 2D band. As expected, the spectra of all
observed materials are quite similar because the main constituent of the doped inks was
graphite. Nevertheless, by analysing the intensity peak ratio between the D and G bands (ID/IG),
we are able to evaluate the level of disorder or defects within the carbon material. The most
common information extracted from the direct comparison between Raman spectrum is that an
Chapter 4
68
increasing tendency of the ID/IG intensity ratios reflects the presence of additional structural
disorder. Here, the ID/IG ratios of the graphite ink alone, with PEDOT, MWCNT and COOH-
MWCNT were 0.30, 0.17, 0.40 and 0.84, respectively. An interesting work have demonstrated a
correlation between the number of edge plane type defects on graphite-based substrates and
the voltammetry of guanine oxidation. The authors have concluded that as the number of
defects sites on the electrode surface increases, there is a small shift in peak potential and,
more importantly, the peak height is enhanced [266]. In our case, it was interesting to observe
that the carbon-based material with less level of disorder (less ratio ID/IG, 0.17) was achieved for
the ink doped with PEDOT nanoparticles, which enabled the higher intensity current for DPV
signaling of 8-OHdG oxidation. This is consistent with the fact that PEDOT displays a structure
where carbon atoms hold an sp2 hybridization. In addition, the material assigned with the higher
number of structural defects (high ratio ID/IG, 0.84) was found to give the weaker DPV response.
4.3.4 Optimization of DPV experimental conditions
Figure 4.5: Successive differential pulse voltammograms of 0.1 mg/mL 8-OHdG in PBS pH 7.4 recorded
(A) without any application of conditioning potential and (B) with a conditioning potential of +0.20V applied
before each measurement.
Here, the oxidation of 8-OHdG in pH 7.4 PBS was found to be around +0.41 V. Initially,
the effect of performing a pre-conditioning step at a specific potential on the DPV
electrochemical performance was investigated (see Figure 4.5). Our data showed that applying
a conditioning potential of +0.20 V for pre-concentrating of 8-OHdG greatly improved the
stability and reproducibility of the DPV measurement. Therefore, the subsequent DPV
measurements were conducted over a positive potential range, after applying a fixed step
potential for the pre-concentration of 8-OHdG on the surface of the modified electrodes.
Afterwards, the influence of the potential and accumulation time during the pre-concentration
step on the electrochemical response of 8-OHdG was also investigated, as shown in Figure 4.6.
Two different potentials were tested and the highest current signal was achieved at +0.20 V so,
this potential value was chosen for further studies. Concerning the accumulation time, it was
Chapter 4
69
observed that the oxidation current increased from 180 s to 300 s and then it decreased for 500
s. Hence, the optimal time pre-concentration for the determination of 8-OHdG by DPV analysis
was chosen to be 300 s.
Figure 4.6: Dependence of the sensor response on the (A) pre-accumulation potential and (B) time of
accumulation during 8-OHdG oxidation in PBS pH 7.4.
Nowadays, one of the most popular electrochemical transducer in biosensor field is carbon-
based surface (graphite, glassy carbon (GC), carbon paste). Due to their great stability and
suitable electronic properties, these nanostructured materials have been widely applied,
nevertheless, appropriate cleaning methodologies are still needed in order to enhance the
material response. Herein, we have studied the effect of different kinds of washing solutions,
such as, ethanol, buffers and acids, on the electrochemical response of our target molecule
(data not shown). We have concluded that the best performance was achieved after applying
voltammetric sweeps on PBS at a pH 7.4 (see Figure 4.7A). More interestingly, by applying the
same cleaning protocol after each DPV measurement (see Figure 4.7B), we were able to
completely regenerate the sensor by returning the current to its original level, enabling multiple
use of the same electrode.
Figure 4.7: (A) Cleaning effect (after CV in PBS pH 7.4) on the 8-OHdG detection by DPV signal and (B)
sensor regeneration after voltammetric cycles performed in PBS pH 7.4.
Chapter 4
70
We have tested the performance of the DPV signal for different pH environments (see Figure
4.8) and, as expected, we have found that the electrochemical oxidation of 8-OHdG is pH
dependent. Our data seems to be in agreement with previous studies in which the anodic peak
potential shifted towards negative values with the increase of the supporting electrolyte pH from
3.0 to 9.0, indicating that the electrochemical process of 8-OHdG is associated with a proton-
transfer process [153]. As shown here, the tendency seems to be that the oxidation peak
current decreases with the increasing of pH (Tris pH 9.5), has it was observed previously for
guanine and adenine oxidation [150]. Although the maximum value of the current intensity was
obtained for PBS when the pH was 6.0, it was quite similar when compared with the pH of 7.4.
Thus, the physiological pH was selected and used in further studies.
Figure 4.8: Differential pulse voltammograms recorded for 8-OHdG solutions prepared in different buffer
solutions, with different pH values.
4.3.5 Analytical applications
4.3.5.1 Calibration curve
Figure 4.9A shows the corresponding dependence of the oxidation peak current on the
concentration of 8-OHdG. As it can be seen, with the increase of concentration, the DPV
signaling was also increased. Figure 4.9B displays the calibration graph for 8-OHdG showing a
linear relationship between the log peak intensity and the log concentration of 8-OHdG over the
range 50-1000 ng/mL. The calibration plot showed excellent linearity (r2 = 0.9919) and the LOD
was found to be 14.4 ng/mL, calculated by the intersection of two straight lines, between the
linear response range and the lower concentration range.
Chapter 4
71
The reproducibility of the modified-electrodes for quantifying 8-OHdG was investigated over the
entire linear range and data showed that the RSD was less than 3.5 %, for five independent
experiments.
Figure 4.9: A) Differential pulse voltammograms for different concentrations of 8-OHdG prepared in PBS
pH 7.4 and (B) calibration plot of the concentration of 8-OHdG.
Although the LOD of our paper-based sensor might not fully satisfy the cut-off of all the
biological samples, it has a relatively wide linear range and a good sensitivity compared to other
works. Moreover, despite the fact that there are some known methodologies for 8-OHdG
assessment down to picoMolar [267], various clinical studies have reported that the levels of
serum 8-OHdG in healthy subjects can suffer variations between 3 and 160 ng/mL
[29],[62]
,[268]
,[269]
,[270]. Thus, our electrochemical label-free sensor is a valuable tool in
providing quick and low-cost information concerning the level of OS at DNA context.
4.3.5.2 Selectivity
The selectivity of the proposed sensor is crucial to grant its successful application.
Meanwhile, one of the main limitations of most conventional methodologies involving biological
samples is the need to include pre-treatment steps in order to minimize possible matrix
interference effects. Herein, the target is to allow a direct sample analysis, with only dilution, if
necessary.
Chapter 4
72
Figure 4.10: (A) DPV recordings for individual solutions with concentrations of 0.1 mM of 8-OHdG,
ascorbic acid and uric acid in PBS at pH of 7.4; (B) DPV recording of a mixture with all of the 3
compounds, in the same concentrations.
Uric acid and ascorbic acid were chosen as main interfering species in the determination of 8-
OHdG, due to their similar structure and high abundance in biological samples, respectively.
Both molecules are also electro-active and, very often, they are oxidized almost at the same
potential value, resulting in the overlap of voltammetric signaling. In order to test the selectivity
behaviour of our sensor, we have investigated the oxidation peak potential of both target-
molecule and interfering species, individually (see Figure 4.10A) and in an equimolar mixture
(see Figure 4.10B). Our data showed that over the tested potential range, ascorbic acid did not
exhibit any redox peak that could affect the analysis (see Figure 4.10A, black line). On the other
hand, an oxidation peak was observed for the uric acid molecule (see Figure 4.10A, dashed
dark blue line), but it occurs at a lower potential value compared with the oxidation potential of
8-OHdG (see Figure 4.10A, light blue line). Meanwhile, both molecules were analysed at the
same concentration level (0.1 mM), however 8-OHdG presented an oxidation peak with higher
current than uric acid. The differential pulse voltammogram obtained from the equimolar mixture
of the three molecules showed that the oxidation potential peak of 8-OHdG in the mixture was
shifted to values above +0.5V may be due to the presence of high concentration of uric acid.
Concerning the peak intensity current, the same value was obtained in comparison with the
individual solution. In addition, the oxidation peak of uric acid was not visible but instead a small
shoulder appeared at lower potentials, enabling a clear separation between the oxidation
process of 8-OHdG and uric acid.
Chapter 4
73
4.3.5.3 Serum samples
Figure 4.11: Differential pulse voltammograms for serum samples diluted 1:10 in different buffers, such as,
(A) Tris pH 9.1, (B) PBS pH 7.4 and (C) Acetate pH 5.1, doped with 1 ug/mL of 8-OHdG.
In order to investigate the applicability of our paper-modified electrode for assessment of
8-OHdG levels in POC, we performed some electrochemical studies using real serum samples.
Biological fluids, such as, urine, blood and serum are complex matrices, composed by high
levels of small biomolecules, some of which are electrochemically active and thereby act as
possible interfering species. Thus, to optimize some experimental conditions, we have tested
the dependence of the oxidation peak potential of both the target-molecule and possible
interferent species for different electrolytes and pH values, aiming for a good voltammetric
separation of the compounds. Figure 4.11 shows the differential pulse voltammograms obtained
for 8-OHdG spiked serum samples 1:10 diluted in different buffers at different pH environments
(basic, neutral and acid). Although it is not displayed here, all the tested diluted serum samples
presented a broad DPV signal between +0.35 V and +0.45 V due to the presence of other
interferent species holding electroactivity characteristics, such as, ascorbic acid and uric acid,
among others. Our data showed that by doping these serum samples with a known
concentration of 8-OHdG we were able to obtain a second DPV peak assigned to the oxidation
of our target-molecule, with the exception of the buffer Tris pH 9.1 that presented no extra peak
(see Figure 4.11A). This outcome is in agreement with our previous pH study showing that the
oxidation peak current is negatively affected by the increasing of pH. For the other tested pH
0.2A
A)
B)
C)
8-OHdG
8-OHdG
E / V vs. Ag/AgClE (V) vs. Ag/AgCl
Chapter 4
74
values (see Figures 4.11B and 4.11C), the best enhancement in the oxidation current of 8-
OHdG was achieved for the sample diluted in PBS pH 7.4. The obtained results are in good
agreement with other similar studies that showed that the peak current of 8-oxoguanine, in the
presence of interferents, was maximum in the pH range 6-8 [271]. In sum, PBS at 7.4 pH was
the optimal choice for 8-OHdG electrochemical detection since the peak current is higher and
sufficiently good separation between peaks was reached.
Figure 4.12: Calibration curve of the concentration of 8-OHdG in diluted serum samples.
Meanwhile, the application of the proposed sensor in the quantification of 8-OHdG was tested in
serum samples by using a spiking approach. All samples were previously diluted 1:10 in PBS
pH 7.4. Three independent measurements were performed for each concentration. Figure 4.12
presents the calibration graph for 8-OHdG concentration obtained from the DPV response. A
linear regression equation (log current (A) = 0.67 log concentration (ng/mL) - 9.3655) was
achieved over the concentration range 10-1000 ng/mL with a correlation coefficient of 0.9958.
Moreover, the reproducibility of the modified-electrodes was confirmed and RSD was less than
2.5 %.
Although other works have been reported with better sensitivity (see Table 4.1), they have used
more complex and expensive protocols in comparison with our paper-based sensor. For
instance, some interesting studies related to the modification of GCE surface with materials that
effectively increased their surface area enabled the detection of 8-OHdG molecule at very low
levels (0.1-0.2 ng/mL) but in both cases the immobilization of antibodies was performed
[146],[252]. One of the main advantages of our 3-electrode system assembled on paper
substrate is their ability to work with small sample volumes. Moreover, there are very few papers
concerning the application of rapid and cost-effective biosensing devices for real serum
samples. Thus, the analysis of 8-OHdG in serum samples was successfully achieved in less
time and with low cost input, allowing the detection of OS biomarkers in biological samples.
Chapter 4
75
4.4 CONCLUSIONS
We have investigated the electrochemical performance of 8-OHdG biomarker at the
surface of paper-modified electrodes for in-situ detection purposes. Different carbon-based
nanomaterials were tested for coating the paper support and it was found that the presence of
the conducting polymer PEDOT could effectively enhance the oxidation peak current of 8-
OHdG. In parallel, several experimental conditions, such as, scan-rate, potential of pre-
accumulation, accumulation time, supporting electrolyte and pH have been carefully optimized
and the electrochemical response of the designed sensor was investigated by means of DPV.
Overall, the combination of electrical properties of PEDOT with the electrochemical sensing of
8-OHdG enhanced the electro-catalytic activity of the working electrode, which resulted in
favourable analytical features, such as, high reproducibility and selectivity, with low cost
resources.
One main advantage of this sensor is easy and quick way to regenerate it, simply by performing
voltammetric cycles in buffer solution, enabling continuous real-time detection of several
samples. The developed electrochemical paper-based sensor showed high sensitivity towards
8-OHdG over the concentration range 50-1000 ng/mL, enabling low detection limit. Thus, the
proposed electrochemical sensor holds high selectivity, reproducibility and stability, which
constitutes a promising low-cost approach to be implemented as an easy-to-use protocol for
sensitive detection of 8-OHdG in biological samples.
Chapter 4
76
Table 4.1: Comparison of different electrochemical sensors for determination of 8-OHdG.
Working
electrode/
Substrate
Working electrode/
Modification
Detection
technique
Range of
linear
concentration
dependence
Type of
Sensor
Real
applicatio
n
Referen
ce
2 electrode
system/ CNT
conductive
paper
___
Chrono-
amperometric
(and
colorimetric)
0-530 nM
Immuno-
sensor
Urine
samples
[256]
3 electrode
system/ GCE
electrode
SWCNTs- Nafion
dispersion DPV 0.03-1.25 M
1.25-8.75 M
Urine
samples [257]
3 electrode
system/ GCE
electrode
Graphene- Nafion
film LSV
0.07-3.64 M
3.64-16.24 M
16.24-33.04 M
Urine
samples [251]
3 electrode
system/ GCE
electrode
MWCNT film LSV 80-5000 nM Urine
samples [272]
3 electrode
system/ GCE
electrode
Poly(3-methyl
thiophene) CV
0.700-35.0 M
35.0-70.0 M
Urine
samples [153]
3 electrode
system/ GCE
electrode
Poly(indole-5-
carboxylic acid)
and chitosan
DPV 0.35-35305 nM Immuno-
sensor
Urine
samples [252]
3 electrode
system/ Pt
electrode
Carbon quantum
dot coated with
Au/SiO2 core-shell
nanoparticles
ECL 0.71-706 nM Immuno-
sensor
Milk
samples [146]
3 electrode
system/ Au
electrode
Phenol polymer EIS 0.35-353 pM MIP Urine
samples [267]
3 electrode
system/ GCE
electrode
Sulfur-doped
graphene DPV 0.002-20 M
Urine
samples [258]
3 electrode
system/ GCE
electrode
DNA
functionalized
graphene
nanosheets
CV
0.005-1.155 M
1.155-11.65 M
11.65-36.15 M
Urine
samples [249]
Au: gold; CNT: carbon nanotube; CV: cyclic voltammetry; DPV: differential pulse voltammetry; ECL:
electrochemiluminescence; EIS: electrochemical impedance spectroscopy; GCE: glass carbon electrode;
HPLC-ECD: High Performance Liquid Chromatography with Electrochemical Detection; LSV: linear sweep
voltammetry; MIP: molecularly imprinting polymer; MWCNT: multi walled carbon nanotubes; PAMAM:
poly(amidoamine); Pt: platinum; SAM: self-assembled monolayers; SWCNTs: single walled carbon
nanotubes.
Chapter 5
77
CHAPTER 5
5 Novel wax-printed paper-based device for a direct
electrochemical detection of 3-nitrotyrosine
The results presented in the chapter were published in Gabriela V. Martins, Ana C. Marques,
Elvira Fortunato, M. Goreti F. Sales, "Wax-printed paper-based device for direct electrochemical
detection of 3-nitrotyrosine", Electrochimica Acta (2018), 284, p. 60-68.
doi: 10.1016/j.electacta.2018.07.150.
Chapter 5
78
5.1 INTRODUCTION
In the near future, clinical analysis are no longer confined to clinical laboratories but
instead, the new trend is to be routinely carried out in other environments, like for instance,
hospital for POC assessment and monitoring. In this context, paper seems the perfect choice as
support material in diagnostic devices, due to its special features, like porosity, affordability,
flexibility, high surface area, sustainability and biocompatibility [171][166][273]. Arduinin and
colleagues have developed a paper-based electrochemical biosensor to detect ethanol in
commercial beers and alcohol oxidase was immobilized onto the modified SPEs [274].
Interestingly, the same group have reported the first example of an integrated paper-based SPE
for the assessment of nerve agents [275]. Looking forward to the last trends in the field of
printed sensors, the investigation of innovative materials is driven by the ability to accomplish
cost-effective devices for continuous monitoring. Printed platforms are becoming a more
affordable and reliable choice in comparison with the conventional three electrode system
because they are disposable, easy-to-use and environmentally-friendly. This issue has been
addressed in a recent paper that developed screen-printed conductive patterns from a water-
based conductive ink assembled on paper substrate [136].
Herein, we have presented the first paper-based electrochemical sensor towards the detection
of 3-NT, a relevant OS biomarker. Mostly due to their high sensitivity and portability,
electrochemical technology has showed to be an attractive approach to detect and quantify
various biological compounds [276][277], in the POC context. Few studies on the
electrochemical behaviour of 3-NT (and their derivatives) at solid substrates, mostly in 3-
electrode system, have been performed [278][279]. The investigation of the redox behaviour of
these electro-active compounds not only contributes to obtain valuable information related to
the molecules, but also enables additional selectivity as different electro-active molecules can
be oxidized and/or reduced at different potential values. In this way, the nature of the conductive
surface of the sensor can be finely tuned and designed towards the detection and quantification
of the target-molecule, hindering the interference of other electro-active compounds [251].
The design and assembly of this voltammetric sensor allowed performing the analysis without
the need for an extra electrocatalyst mediator specie. This innovative sensor was designed by,
first, modifying paper to become a hydrophobic support and, second applying carbon and silver
conductive inks to generate a three electrode-system on a small spot. In addition, we have used
the advantage of paper to design a low-cost, disposable and user-friendly flexible platform for
POC assessment. The optimization performed on the fabrication of the device in parallel with
the experimental conditions of the electrical measurements allowed to obtain an integrated 3-
electrode device that sensitively and selectively quantified 3-NT biomarker (see Figure 5.1).
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Figure 5.1: Schematic illustration of the different steps related to the sensor device, namely, A) the
electrochemical apparatus for biological samples assessment; B) photo and morphological
characterization of the paper-based electrodes; and C) the assembly of the electrochemical sensing
platform.
5.2 EXPERIMENTAL SECTION
5.2.1 Reagents and Materials
All reagents were of analytical grade and used as supplied without purification. Potassium
hexacyanoferrate III (K3[Fe(CN)6]), potassium hexacyanoferrate II (K4[Fe(CN)6]) trihydrate,
dipotassium hydrogen phosphate (K2HPO4) and L-ascorbic acid p.a. were obtained from Riedel-
de-Häen; hexaammineruthenium (III) chloride (Ru(NH3)6Cl3) from Acros Organics; potassium
chloride (KCl) from Merck; PBS tablets from Amresco; potassium di-hydrogenophosphate
(KH2PO4) from Panreac; TRIS (hydroxymethyl)aminomethane (TRIS) from Fisher BioReagents;
3-NT (98% purity) from Alfa-Aesar; L-tyrosine pure from AppliChem; uric acid from Sigma-
Aldrich and creatinine from Fluka.
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The supporting electrolytes and the buffer solutions were prepared with ultrapure water Milli-Q
laboratory grade. The pH measurements were performed with a pH meter, from Crison
Instruments, GLP 21 model. All experiments were carried out at room temperature.
5.2.2 Fabrication of paper-based SPE
Herein, white office paper (Bright White Paper for colour laser printing 160 g/m2) was
employed as the substrate for the fabrication of the paper-based SPE (see Figure 5.2). A vector
drawing software (Adobe Illustrator from Adobe Systems Software) was used to design the
electrodes configuration and, afterwards, the hydrophobic pattern was printed directly onto the
paper by means of a wax printer (Xerox ColorQube model 8580). Then, the wax printed paper
was cured at 120 °C for 3 min, allowing the wax to diffuse vertically throughout the paper. After
producing the hydrophobic confinement, the 3-electrode system was manually painted using a
silver ink (AG-530 Flexible Silver from Applied Ink Solutions) as the pseudo-reference electrode
and a carbon-based ink (PE-C-774 Carbon Resistive Ink from Applied Ink Solutions) to print
both working and counter electrodes. In order to apply the three-electrode system we have used
tape-masks designed (kapton tape, 0.25 mm thickness), which were patterned by means of a
laser-cutting machine (model VLS 3.50, CO2 infrared laser with 10.6 m wavelength from
Universal Laser Systems).
Figure 5.2: Detailed scheme of the fabrication of the paper-based electrodes.
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5.2.3 Electrochemical assay
Electrochemical measurements were conducted by using the fabricated paper-based
electrodes composed by a carbon-ink electrode as the counter and working (5 mm diameter)
electrodes. Silver-ink was used as pseudo-reference. For comparison studies, commercial
carbon-screen printed electrodes (SPEs-DRP-C110) from DropSens were also used. The
electrodes were electrochemically characterised with a potentiostat/galvanostat from Metrohm
Autolab and a PGSTAT302N with a FRA module, through an interface switch box from BioTid
Eletrónica and controlled by ANOVA software. All the presented potential values are related to
the silver pseudo-reference of the SPEs.
Initially, the working electrode was electrochemically cleaned by performing voltammetric
sweeps between -0.20 V and +1.50 V in PBS pH 7.4, until a stable voltammogram was
obtained. Before use, sensors were air dried and stored at room temperature. [Fe(CN)6]4-/3-
couple and [Ru(NH3)6]3+
were selected as electrochemical indicators to evaluate the
performance of the paper-based SPEs by CV. The measurements were performed by covering
the three electrodes with 200 l of an appropriate solution. Two different supporting electrolytes
were tested during this study, namely, 0.1 M KCl and PBS pH 7.4. For the [Fe(CN)6]4-/3-
probe,
CV was performed from -0.50 V to +0.70 V and in the case of the [Ru(NH3)6]3+
probe, a potential
range between -0.70 V and +0.20 V was chosen.
5.2.4 Surface characterization of the paper-based SPEs
SEM analysis was performed in order to evaluate the morphology of the surface of the
fabricated paper-based SPEs. The samples were examined in a Carl Zeiss AURIGA
Crossbeam SEM-FIB workstation, by using an accelerating voltage of 5 kV with an aperture
size of 30 mm. Cross section imaging was also performed to estimate the thickness of the
carbon layer on the working electrode.
5.2.5 Detection of 3-Nitrotyrosine onto the paper-based SPE
The electrochemical response of 3-NT was firstly investigated by performing CV
recordings between -1.0 V and +1.0 V and, afterwards, by means of SWV, without the use of
any redox probe. The voltammetric parameters used during SWV were pulse: amplitude 20 mV,
frequency 5 Hz and scan-rate 25 mV/s. Electroactivity studies were made in different buffer and
supporting electrolytes, such as, 10 mM PBS pH 7.4, 0.1 M Phosphate Buffer pH 6-8, 1 M TRIS
buffer pH 8.3 and 0.1 M KCl pH 5.8. Calibration curves were performed with 3-NT standard
solutions ranging from 250 nM and 1 mM, prepared in the appropriate solution.
5.2.6 Selectivity assay
Under this study, tyrosine, ascorbic acid, uric acid and creatinine were chosen as
interfering molecules. The selectivity studies were carried out, firstly, by performing SWV
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measurements for each individual compound at high concentration, in the conditions described
above. Afterwards, 3-NT calibration curves were carried out with and without the presence of
the interfering tyrosine molecule.
5.3 RESULTS AND DISCUSSION
5.3.1 Electrochemical performance of the paper-based SPEs
Herein, the electrochemical characterization of the paper-based electrodes was carried
out with CV method by choosing [Fe(CN)6]4-/3-
couple as the redox-probe. Figure 5.3 displays
the CV recordings at several scan-rates, in 2 different electrolyte solutions, namely, 0.1 M KCl
(Figure 5.3A) and PBS pH 7.4 (Figure 5.3B). All experiments were performed in triplicate.
Figure 5.3: Cyclic voltammograms for 5 mM [Fe(CN)6]4-/3-
redox couple in A) 0.1 M KCl solution and B)
PBS pH 7.4, at different scan-rates; Plot representation of both the anodic and cathodic peak currents
versus the square-root of the scan-rate for 5 mM [Fe(CN)6]4-/3-
redox couple in C) 0.1 M KCl solution and
D) PBS pH 7.4.
The analysis of the results showed that both electrolytes enabled good anodic and cathodic
peaks with high peak amplitudes. Moreover, the increasing separation between the reduction
and oxidation peaks potentials with increasing scan-rate reflects the quasi-reversible redox
process of [Fe(CN)6]4-/3-
redox couple at the surface of our electrodes. Therefore, the currents of
both anodic and cathodic peaks were plotted against the square-root of the scan-rate for the 2
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studied systems, 0.1 M KCl and PBS, as can be seen in Figure 5.3C and Figure 5.3D,
respectively. Our data showed a linear dependent behaviour along the scan-rate range 5 - 400
mV/s which is an indication of a diffusion-controlled mechanism responsible for the
electrochemical process occurring at the electrode surface [280].
As previously explained, the nature of the conductive ink and its interaction with the substrate
are key elements for the electrochemical performance of the electrode. Therefore, in most
cases, the real area where electron exchange takes place, also known by active
electrochemical surface area, is different from the geometrical surface of the electrode. To
overcome this issue, the effective active surface area of the electrode can be estimated using
the Randles-Sevcik equation [281]:
Ip = 268600 x n3/2
x A x D1/2
x C x ʋ1/2
where Ip is the peak current intensity (A), n is the number of electrons transferred in the
electrochemical reaction, A is the electrode area (cm2), D is the diffusion coefficient of the
analyte, C is the bulk concentration of the analyte (mol/cm3) and ʋ is the scan-rate (V/s).
Herein, we have employed the data from 5 mM [Fe(CN)6]4-/3-
redox couple in 0.1 M KCl since it
was the system with better electrochemical reversibility behaviour. Accordingly, from the graphic
representation of the peak current versus the square-root of the scan-rate we have applied the
slope of this plot to the previous equation in order to estimate the electroactive surface area of
our fabricated electrodes. Under these conditions, the value obtained for the effective active
surface area of the electrodes was 0.138 ± 0.007 cm2. Although this result is lower than the
respective geometric area of the electrode, the activity ratio Sa/Sg, where Sa is the electroactive
surface area and Sg is the geometric surface area, is quite similar to the respective value
obtained for the commercial carbon-SPEs from Dropsens, which was around 30% (see Figure
5.4).
Figure 5.4: A) Cyclic voltammograms for 5 mM [Fe(CN)6]4-/3-
redox couple in 0.1 M KCl, at different scan-
rates and B) plot representation of both the anodic and cathodic peak currents versus the square-root of
the scan-rate for 5 mM [Fe(CN)6]4-/3-
redox couple.
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In this study, CV was employed to investigate the electrochemical performance of the carbon-
based electrode towards the negatively charged [Fe(CN)6]4-/3-
redox couple and the positively
charged [Ru(NH3)6]3+
. Herein, [Ru(NH3)6]3+
was selected as a redox probe because it undergoes
a rapid and chemically quasi-reversible one-electron redox reaction, with both the oxidized and
reduced forms of the ruthenium complex. Figure 5.5A shows the variation E for both type of
probes versus the scan-rate, in 0.1 M KCl. It was interesting to observe a quite similar
behaviour for [Fe(CN)6]4-/3-
and [Ru(NH3)6]3+
probes, that was a high increase of E for lower
scan-rates and at higher scan-rates it tends to stabilize. Moreover, the ratio IpA/IpC was
calculated and plotted over the scan-rate range 5 - 400 mV/s, for both probes (Figure 5.5B).
The estimation of this parameter allowed evaluating the reversibility character of the
electrochemical system. As this ratio moves way from the value of 1, the system becomes less
reversible [125]. The analysis of both graphs (E and IpA/IpC versus scan-rate) seems to
indicate a higher reversibility for [Fe(CN)6]4-/3-
system because peak-to-peak separation is lower
and the ratio IpA/IpC is closer to 1 in comparison with [Ru(NH3)6]3+
system. However, this may be
related to the fact that only one ionic form of ruthenium was present, against the redox couple of
iron.
Figure 5.5: Effect of the different redox probes upon the electrochemical response. A) plots the peak
potential separation (ΔE) versus the scan-rate and B) the anodic and cathodic peak current ratio (IpA/IpC)
versus the scan-rate for [Fe(CN)6]4-/3-
and [Ru(NH3)6]3+
probes at 5 mM concentration in 0.1 M KCl; Plots of
current peak versus the probe concentration for C) [Fe(CN)6]4-/3-
and D) [Ru(NH3)6]3+
, at a scan-rate of 50
mV/s, in 0.1 M KCl and PBS pH 7.4.
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Different concentrations of the redox probe, for an applied potential scan-rate of 50 mV/s, were
tested and the current-voltage characteristics were recorded for [Fe(CN)6]4-/3-
(Figure 5.5C) and
[Ru(NH3)6]3+
(Figure 5.5D), in 0.1 M KCl and PBS pH 7.4 conditions. Although in both cases, the
current peak amplitude increases with increasing probe concentration, the tendency of the
response is different for the 2 types of probe. Not only the current peak amplitude is higher for
the [Fe(CN)6]4-/3-
couple, the [Ru(NH3)6]3+
probe did not seem to be sensitive to a supporting
electrolyte variation, as depicted in Figure 5.5D, from the identical tendency in PBS and KCl
solutions. In sum, for future applications, [Fe(CN)6]4-/3-
redox-probe seems to be a more suitable
choice.
5.3.2 Morphological characterization of the paper-based SPEs
SEM analysis was also conducted to further investigate the surface morphology of the
carbon electrode assembled onto the paper substrate. As can be seen in Figures 5.6A-B, the
WE surface of carbon ink is a more or less uniform carbon layer composed by small islets, with
very few cracks on the surface. In addition, the paper-based electrodes were vertically cut, and
cross-section imaging were performed. Figure 5.6C shows the thickness of the applied carbon
layer: by performing suitable measurements, the obtained values ranged from 33.1 to 39.9 m.
Furthermore, as depicted in Figure 5.6D, it was quite obvious to distinguish the borderline
between the applied carbon layer and the paper substrate with its characteristic cellulose fibers.
Figure 5.6: SEM images of the A) and B) WE carbon-surface at different magnifications and C) and D)
Cross-section imaging of the carbon-layer at different magnifications.
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5.3.3 Direct detection of 3-Nitrotyrosine
The electrochemical behaviour of 3-NT was studied by performing CV swept over the
potential range -1.0 V and +1.0 V, in phosphate buffer solution. Figure 5.7 displays two well-
defined oxidation peaks at +0.10 V and +0.80 V and, in the negative-moving direction, we were
able to observe a cathodic peak at -0.84 V. Although not displayed here, during our study we
found that the first anodic peak at +0.10 V was dependent of the occurrence of the cathodic
peak, meaning that this oxidation reaction corresponds to the species formed at the carbon-
based surface after 3-NT reduction at -0.84 V. According to previous studies [282][131][209],
this reduction peak was assigned to the four-electron reduction of the nitro group to the
corresponding hydroxylamine. Although mechanistic studies related to the electrochemical
behaviour of 3-NT are not very common, herein, the characteristic oxidation peak of L-tyrosine
structure was +0.80 V, attributed to a two-electron and two-proton process [283][279]. So, for
further electrochemical experiments, this potential peak was chosen to study the electro-
oxidation of 3-NT.
Figure 5.7: CV recordings over the potential range -1 V to +1 V in 0.1 M phosphate buffer with (colour line)
and without (dashed line) 1 mM of 3-NT, at a scan-rate of 50 mV/s, and in the inset figure the chemical
structure of 3-NT.
Under the scope of this work, SWV was the preferred voltammetric technique to follow the
oxidation behaviour of 3-NT due to its high sensitivity combined with a fast response [113].
Firstly, as illustrated in Figure 5.8A, different electrolytes at different concentrations, such as,
0.1 M phosphate buffer, 10 mM PBS, 1 M TRis buffer and 0.1 M KCl were investigated to
determine the most appropriate supporting electrolyte matrix. Although there were substantial
differences related to the oxidation peak current, mostly due to the salts concentration, the more
interesting issue here was the shift in the peak potential. Moreover, we observed that 3-NT in
KCl solution (at a 5.7 pH) did not exhibit any peak over the potential range +0.60 V to +1.0 V.
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So, along this study, phosphate buffer was chosen as the supporting electrolyte, showing the
best peak amplitude/resolution and an oxidation peak displaced at less positive value, thereby
revealing a more favourable process.
Afterwards, the effect of pH on the electrochemical oxidation of 3-NT was also studied by SWV
in a 0.1 M phosphate buffer solution containing 1 mM of 3-NT (Figure 5.8B1). Our data showed
that with increasing pH from 6 to 7.5, the oxidation potential was moved towards less positive
values and, in addition, a linear-dependent behaviour on the pH value of the buffer solution was
found (Figure 5.8B2), implying that both electrons and protons are involved in the electro-
oxidation of 3-NT [207]. Furthermore, results showed that the oxidation peak current changed
slightly with pH variation. Thus, subsequent experiments were carried out at pH 7.4, not only
because it was near the best amplitude obtained, but also because we target to work under
physiological conditions. In sum, besides electrolyte composition, the pH value also influences
the detection performance of 3-NT.
Figure 5.8: SWV response of 1 mg/mL 3-NT A) in different supporting electrolyte solutions and B1) in 0.1
M phosphate buffer solution at different pH values ranging from 6 to 8. B2) Plot of the potential value of the
SWV versus the pH obtained in 1 mg/mL 3-NT in phosphate buffer solution.
5.3.4 Calibration and interference assay
Under the optimized parameters, the carbon-based electrode was tested with different
concentrations of 3-NT. As displayed in Figure 5.9A, the peak current in SWV increased for
increasing concentrations. Figure 5.9B shows the calibration curve for 3-NT with and without the
application of pre-accumulation potential [+0.40 V] previous to the reading. The use of a pre-
accumulation step has already been investigated as an approach to enhance the
electrochemical signal for direct detection of electro-active species [257]. Herein, both results
presented a quite similar behaviour and the sensor device showed a good linearity (r2 = 0.9937
and 0.9927) over the concentration range 500 nM to 1 mM. The LOD, calculated by applying 3σ
of the blank as three times the standard deviation of the first 3-NT concentration that enabled a
current measurement, was found equal to 49.2 nM. Also, for 3 different independent
experiments, the paper-modified electrodes showed excellent reproducibility with RSD lower
than 1.5 %.
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Figure 5.9: A) SWV recordings of 3-NT at different concentrations, in 0.1 M phosphate buffer at 7.4 pH
(inset figure is for lower concentrations) and B) calibration curve of 3-NT, with and without the application
of an accumulation potential.
Although this LOD was found slightly higher in comparison with other electrochemical studies
reported in literature [209][278][47], our work appears as the voltammetric sensor device with
more potential to become portable and suitable for POC application. In addition, some of these
approaches have incorporated nanostructured materials during the sensor assembly as a
requirement to obtain the required low detection limits. Firstly, the introduction of a flexible and
low-cost substrate as paper to be use as the sensing platform enables a new generation of
sensors towards the future sustainability. Moreover, the integration of the 3-electrodes system
in the same structure holds many advantages, such as, simplicity, quick response, small feature
size and low-cost instrumentation. Another issue that still needs to be well understood is the
clinical value of the basal levels of 3-NT reported. In spite of the fact that has been an
increasing evidence that the plasma concentrations of free 3-NT in healthy humans are in the
order of 1 nM [29], another work have reported basal concentrations for 3-NT in a certain
biological matrix that varied up to a factor of 1000 between different methods [44]. Thus,
differences in the quality of the samples or in the composition of the study groups may
contribute to a variability in 3-NT concentration.
Figure 5.10: A) Electrochemical response of tyrosine, ascorbic acid, uric acid and creatinine over the
studied potential range and B) the curves of calibration for 3-NT only and in the presence of 10 μM of
tyrosine.
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Herein, tyrosine, ascorbic acid, uric acid and creatinine were chosen as interfering molecules
due to their similar structure, their presence in biological fluids and also their ability to behave
like electro-active species. In this work, we have performed SWV recordings with each
individual molecule in phosphate buffer solution under the potential range applied to the
detection of 3-NT. As can be displayed in Figure 5.10A, only tyrosine gave a signal around
+0.67 V, while the other molecules did not show any type of electro-activity. Since this oxidation
potential is earlier to the oxidation of 3-NT, we applied the pre-accumulation potential described
before (+0.40 V) to eliminate the interference of tyrosine. The results are illustrated in Figure
5.10B and there seems to be a strong evidence that, in the presence of a high concentration of
tyrosine, the calibration curve of 3-NT holds the same tendency with a slight variation in the
slope. Thus, in sum, we have demonstrated the fabrication and application of a label-free
sensor device towards the detection and quantification of 3-NT. Although the desirable LOD
was not yet achieved, in the future, to enhance the performance of this sensor some
amplification steps can be included, such as, the introduction of nanomaterials [107] or even the
integration of molecular imprinting materials [284]. In addition, the previous approaches can
also improve the selectivity behaviour of this electrochemical sensor when applied to complex
biological matrices.
5.4 CONCLUSIONS
In the present work we have designed and fabricated a paper-based device for
electrochemical detection of 3-NT. Herein, these carbon-modified electrodes were
electrochemically characterized, and the outcome data was quite similar to the available
commercial-ones, which are constructed in a ceramic based support. Afterwards, the electro-
oxidation of 3-NT was assessed over the concentration range 500 nM to 1 mM. The
optimization of some experimental parameters, such as, pH and supporting electrolyte solution
was performed and enabled a low detection limit for the direct determination of 3-nitrotyrosine.
As a proof-of-concept, the proposed electrochemical label-free sensor holds high sensitivity,
selectivity and reproducibility, representing a valuable alternative to the complex, time-
consuming and high costly methodologies currently used. In addition, the fabrication of sensing
devices on flexible substrates is likely to find further application in-loco and routine analysis
outside the laboratory environment.
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Chapter 6
91
CHAPTER 6
6 Electrochemical paper-based biosensor for label-
free detection of 3-nitrotyrosine in human urine
samples using molecular imprinted polymer
The results presented in this chapter are currently submitted as Gabriela V. Martins, Ana C.
Marques, Elvira Fortunato, M. Goreti F. Sales, "Electrochemical paper-based biosensor for
label-free detection of 3-nitrotyrosine in human urine samples using molecular imprinted
polymer" (2018).
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6.1 INTRODUCTION
As mentioned previously, 3-NT is a sub-product generated during protein attack by free
radicals [285][286] and has been proposed as a biomarker of OS [16]. An interesting overview
concerning the routes for nitric oxide metabolism and its role in 3-NT biosynthesis was recently
presented [287]. Aiming for an accurate quantification, different methodologies have been
proposed including electrochemical detection [45][75]. For now, we still do not reach a
consensus related to the basal levels of free 3-NT in human urine samples but most reported
methods have presented concentrations of few nM in healthy subjects. In sum, portable, low-
cost and quick analysis of clinical samples is an urging need in POC screening of 3-NT. To
overcome this issue, biosensor technology has been introduced as the design of analytical
devices that respond to a specific chemical specie in a biological matrix [288]. Among others,
biosensors gained great interest due to their special features, such as, automation, quick
response time, high sensitivity, miniaturisation and portable-size [92].
In parallel, SPE have been preferred against the conventional three-electrode system because
these are disposable, easy-to-use and may operate with minimal volumes of analyte solutions.
Following the last trends, paper has been chosen as the ideal support material to develop a
screen-printed biosensor device for 3-NT, due to their special characteristics, such as, flexibility,
porosity, biocompatibility, facile modification and functionalization, low-cost and sustainability
[164][160]. Recently, a facile paper-based visual sensor for detecting anthrax biomarker has
been developed by using filter paper immobilized with Tb/DPA@SiO2-Eu/GMP, enabling direct
observation of the colour switch from green to red by naked eyes under a UV lamp [289].
Moreover, a disposable paper-based platform with a bipolar electrochemical device was
reported for the sensitive detection of the cancer marker prostate specific antigen (PSA),
highlighting important sensing characteristics, such as simplicity, portability and disposability
[290].
The other important component of an SPE is the biorecognition element. In order to achieve
high sensitivity and selectivity standards, molecular imprinting technique coupled with
electrochemical transduction has proven its potential in the fabrication of innovative biomimetic
sensors. Among the different ways of MIP synthesis, electropolymerization enables the direct
deposition of the sensing material on the transducer surface, allowing a facile and easy control
of the film growth [205]. In terms of transducing element, electrochemical approaches have
been widely applied for the detection and quantification of relevant electro-active markers and,
due to their high sensitivity and selectivity [291][261][292].
Currently, the most commonly studied oxidation biomarkers of damaged DNA are purine bases
(adenine and guanine) and 8-OHdG [30]. In this work, we have taken advantage of paper to
design a low-cost and user-friendly electrochemical sensing platform against 3-NT, a sub-
product originated during protein damage. Although there are few reported sensors for detection
of 3-NT in biological matrices (Table 6.1), generally the use of signal amplification steps, such
Chapter 6
93
as, the incorporation of nano-structured materials, was required in order to obtain low detection
limits.
Table 6.1: Comparison of the different sensors for 3-nitrotyrosine detection in biological matrices.
Method Substrate Linear
Range LOD Sample Ref.
Electrochemical detection
+ MIP
Bimetallic Fe/Pd
nanoparticles 21.6-3833 nM 5.3 nM
Human blood
Urine [209]
Electrochemical detection
+ MIP
GCE electrode
with AuNPs 0.2-50.0 M 50.0 nM
Human serum
Urine [47]
Fluorescent detection +
MIP Carbon dots 0.050-1.85 M 17 nM Human serum [210]
HPLC detection with SPE
+ MIP ___ 11.1-243 nM 3.1 nM Human urine [293]
Electrochemical detection Mercury drop
electrode 6.6-597 nM 0.25 nM
Cerebrospinal
fluid, Plasma [278]
SPR detection Graphene 2.21-4421 pM 0.57 pM Human serum [294]
AuNPs: gold nanoparticles; Fe: iron; GCE: glassy carbon electrode; HPLC: High-performance liquid
chromatography; MIP: molecular imprint; Pd: palladium; SPE: solid phase extraction; SPR: surface
plasmon resonance.
Herein, we have employed a molecular imprinting approach based on the assembly of a non-
conducting phenol matrix in order to develop a sensitive and selective label-free electrochemical
biosensor assembled on a paper substrate and aimed at the sensitive detection of 3-NT
biomarker (see Figure 6.1).
Figure 6.1: Illustration of the sensor film fabrication by molecular imprinting for recognition of 3-
nitrotyrosine.
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6.2 EXPERIMENTAL SECTION
6.2.1 Reagents and Materials
All reagents were of analytical grade and used without further purification. Buffer and
electrolyte solutions were prepared with ultrapure water Milli-Q laboratory grade. The pH
measurements were made in a pH meter from Crison Instruments, GLP21 model. Potassium
hexacyanoferrate III (K3[Fe(CN)6]), potassium hexacyanoferrate II (K4[Fe(CN)6]) trihydrate and
dipotassium hydrogen phosphate (K2HPO4) were obtained from Riedel-de-Haen; potassium
dihydrogenophosphate (KH2PO4) was obtained from Panreac; potassium chloride (KCl) was
from Merck; phenol (C6H5OH, for molecular biology) and sulphuric acid 95-97% (H2SO4) were
from Sigma-Aldrich; PBS tablets from Amresco and 3-NT (98%) from Alfa Aesar. Potassium
phosphate buffer solutions (0.1 M, pH 6.0) and (0.2 M, pH 7.0) were prepared by mixing the
proper amount of K2HPO4 and KH2PO4. All experiments were performed at ambient
temperature.
6.2.2 Apparatus
The biosensor device was assembled on fabricated paper-based electrodes composed
by a carbon-ink as the counter and working (5 mm diameter) electrodes and, silver-ink as
pseudo-reference electrode. Detailed description of the procedure of fabrication of these
electrodes was described elsewhere [295]. The electrochemical measurements were carried out
with a potentiostat/galvanostat from Metrohm Autolab and a PGSTAT302N with an FRA
module, through an interface switch box from BioTid Eletrónica, and controlled by ANOVA
software. All the presented potential values are against the silver pseudo-reference.
Raman spectroscopy was performed by using a Thermo Scientific DXR Raman microscope
system with a 785 nm excitation laser, combined with a 50 objective magnification. Raman
spectra were collected with an incident maximum laser power of 3 mW and through a slit
aperture of 50 μm. Photobleaching was set to 20 min. For Raman measurements, data analysis
was performed with OMNIC software.
SEM analysis was performed in order to evaluate the morphology of the polymeric films
modified on the electrode surface, by using a Carl Zeiss AURIGA Crossbeam SEM-FIB
workstation, operating with a voltage of 5 kV.
6.2.3 Electrochemical assay
Initially, the carbon surface of the electrodes was electrochemically cleaned by
performing voltammetric sweeps between −0.2 V and +1.5 V in PBS at pH 7.4, until a stable
voltammogram was obtained (more or less 100 cycles). Before use, sensors were well dried
with nitrogen and stored at room temperature.
CV assays were performed at different potential windows in order to evaluate the electroactive
behaviour of the compounds. Each modification performed on the paper-based electrodes was
Chapter 6
95
electrochemically characterized by EIS of a 5 mM solution of K3[Fe(CN)6] and K4[Fe(CN)6],
prepared in phosphate buffer solution (0.1 M, pH 6.0). EIS experiments were carried out over
the frequency range 0.01 Hz to 100 kHz. All measurements were performed by covering the
three electrodes of the paper-based sensor with a volume solution of 200 l.
The detection of 3-NT molecule was followed by means of DPV in the potential range 0 V to
+0.4 V, at a scan-rate 25 mV/s, pulse amplitude 25 mV and pulse width 50 ms. Before DPV
measurement, the analyte was pre-concentrated at the electrode surface by applying a potential
value of 1 V, for an optimized time. All experiments were conducted in triplicate at room
temperature. Calibration curves were made with fresh 3-NT standard solutions ranging from 100
nM to 1 mM.
6.2.4 Assembly of the imprinted-based biosensor
The MIP film was deposited on the surface of the carbon-coated electrode through bulk
polymerization. Initially, the electrodes were subjected to 5 voltammetric scans over the
potential range −0.2 V to +1.5 V in H2SO4 0.5 M solution, as an electrochemical cleaning
procedure. Afterwards, the electropolymerization was performed by applying cyclic voltammetry
sweeps ranging between +0.2 V and +0.8 V, at a scan-rate of 50 mV/s, in KCl solution (0.1 M,
pH 5.9) containing both phenol monomer and the template molecule 3-NT. Then, the template
removal was made by incubation of the working electrode in a methanol:water solution (1:10,
v/v) for 2 hours, at room temperature, followed by 1 hour stabilization in phosphate buffer
solution. These solutions were prepared daily and all experiments were carried out at room
temperature. Moreover, excluding the presence of 3-NT molecule, a similar procedure was
taken in the fabrication of blank control material, named NIP-modified electrodes.
6.2.5 Analysis of urine samples
The selectivity features of the MIP-based sensor were directly assessed in human urine
samples, due to their complex matrix. The urine samples were collected in sterile falcon tubes to
avoid contamination and, afterwards, the fresh urine samples were frozen and stored in aliquots
of 1 mL. Before the analysis, the samples were diluted in phosphate buffer solution (0.1 M, pH
6.0) in a 1:10 ratio. The detection of 3-NT in urine samples was performed by immersing the
MIP-based device in the diluted samples, followed by a negative potential accumulation and
finally DPV analysis. The same procedure was taken to the NIP control electrode.
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6.3 DISCUSSION AND RESULTS
6.3.1 Electrochemical study
Figure 6.2: Cyclic voltammograms of 3-nitrotyrosine in A) PBS solution, at different scan directions, over
the potential range -1 V to +1 V; B) three different electrolyte solutions, over the potential range -0.4 V to
+1.2 V; and C) KCl solution, over the potential range +0.2 V to +0.8 V. Cyclic voltammograms of D) phenol
and 3-nitrotyrosine, individually and E) mixture phenol + 3-nitrotyrosine. Chemical representation of F)
phenol and G) 3-nitrotyrosine.
In the last years, the study of electrochemical oxidation of organic substances has been
employed as a promising technique for analytical purposes. Herein, in order to evaluate the
electrochemical behaviour of 3-NT in PBS solution (10 mM, pH 7.4), cyclic voltammetric sweeps
were scanned in different directions over the potential range 1.0 V and +1.0 V, at a scan-rate
of 50 mV/s. As depicted in Figure 6.2A, two well-defined oxidation peaks were observed at
+0.20 V and +0.93 V, and a reduction peak around -0.85 V. This outcome is in accordance with
our previous electrochemical study performed with the same modified-electrodes [295] and
interestingly, we have found that the first anodic peak at +0.20 V depended of the occurrence of
the cathodic peak. Hence, the results obtained by CV seemed to indicate that the earlier
oxidation reaction corresponded to sub-species generated at the carbon surface after the
reduction of 3-NT occurred at 0.85 V.
Although there are already some electrochemical sensor devices based on the electro-oxidation
of 3-NT for detection applications [209][47], to the best of our knowledge, this is the first time
that this low oxidation peak potential (+0.20 V) has been chosen to follow this target molecule.
The main advantages of working with this narrow potential window are that the oxidation
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reaction occurring on the surface of the modified electrode is more favourable; there is no need
to apply high voltage to the electrode system that can cause an overcharge and, consequently,
could compromise the performance of the silver pseudo-reference; and also, it allows to
eliminate the interference of other electroactive species found in biological matrices (uric acid,
ascorbic acid, etc), which occurs at higher potential values [296][251][259].
The effect of different buffer and electrolyte species on the oxidation peak potential was also
assessed by CV in a solution containing 4 mM of 3-NT (Figure 6.2B). For this study, we have
tested phosphate buffer solution (0.2 M, pH 7.0), PBS solution (10 mM, pH 7.4) and KCl
solution (0.1 M, pH 5.9). Despite the difference on the salt species and concentration, it could
be noticed that PBS and phosphate buffer solutions exhibited similar current and potential
behaviour, while in KCl solution we could observe that the oxidation peak potential of 3-NT
shifted to a more positive position. This last result can be advantageous in order to guarantee a
wider potential window in which 3-NT molecule does not display electroactive properties.
Furthermore, Figure 6.2C shows cyclic voltammograms in KCl solution alone and 3-NT in KCl
solution, confirming that in the potential range +0.2 V to +0.8 V there is no evidence of oxidation
reaction of 3-NT.
Under the scope of this work, phenol has been chosen as the structural monomer to grow the
imprinted polymeric matrix. It is a non-conducting material, holding good stability properties and
typically known electro-oxidation properties. Several mechanisms regarding the oxidation of
phenol and phenol derivatives have been described on gold [49] and platinum [50][51] surfaces,
but less on glassy carbon electrodes [52]. Among others, experimental parameters, such as,
temperature, concentration of the monomer, electrolyte type and pH of the applied electrolyte
have proven to influence the electrochemical oxidation of phenol [48][51][54].
Figure 6.2D displays cyclic voltammograms scanned over the potential range +0.2 V to +1.2 V,
at a scan-rate of 50 mV/s, for a 10 mM solution of phenol and a 4 mM solution of 3-NT,
individually, both prepared in KCl 0.1 M solution (pH 5.9). As expected, these results indicated
an irreversible oxidation reaction of the phenol monomer occurring on the electrode surface,
around +0.9 V. Furthermore, in accordance with our previous data, 3-NT alone exhibited an
anodic peak at higher potential values. Meanwhile, Figure 6.2E showed the voltammetric sweep
performed on the mixture phenol and 3-NT in KCl 0.1 M solution and only one well-defined
oxidation peak was observed, meaning that under this experimental conditions it was not
possible to separate accurately phenol (Figure 6.2F) and 3-NT (Figure 6.2G) electro-oxidations.
In sum, the electropolymerization of phenol was carried out herein over the potential range
+0.2 V to +0.8 V, in order to avoid any kind of oxidation reaction that could modify or entrap
irreversible 3-NT molecule in the polymeric matrix.
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6.3.2 Electropolymerization of phenol - MIP versus NIP
Figure 6.3: Cyclic voltammograms of NIP and MIP electrodes during electrochemical polymerization of
phenol for the A) 1st and B) 5
th scan cycle, in KCl solution (0.1 M, pH 5.9). EIS obtained for each step of
the construction for C) NIP and D) MIP electrodes, in 5 mM solution of K3[Fe(CN)6] and K4[Fe(CN)6]
prepared in phosphate buffer solution (0.1 M, pH 6.0).
Herein, electropolymerization was the approach chosen to assemble the biomimetic
material because we can precisely tune the thickness and growth of the film in-situ, without
using additional initiator species. Briefly, electro-oxidation of phenol occurs through the
formation of the phenoxy radical, that can react with other species present in the solution
generating products or else reacts with other phenol molecules producing a dimeric radical.
Usually, the rate-determining step for phenolic compounds is one electron reaction, such as
phenoxy radical forming [47]. The polymerization curves corresponding to the MIP and NIP
imprint polymers for the 1st and 5th scan cycles are shown in Figure 6.3A and Figure 6.3B,
respectively. As expected, voltammograms displayed an irreversible oxidation reaction of the
monomer for both NIP and MIP materials, with NIP holding a higher current peak. It was also
observed that the potential of the oxidation peak did not suffer any change with the increasing of
the number of cycles. This behaviour was attributed to the electrode fouling produced by the
formation of a non-conductive polymeric layer resulting from phenol oxidation that blocks the
electrode surface. Moreover, the current peaks gradually decreased with the number of cycles,
which is another characteristic assigned to the growth of non-conducting polymers [49]. In
addition, no significant difference was observed in the cyclic voltammograms with and without
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the template molecule, shown in Figure 6.3B, by the similar amplitude current obtained in the
end of the polymerization reaction.
One of the great advantages of using an electropolymerization approach is their ability to finely
tune the amount of deposited polymer, by adjusting the duration of the applied voltage (number
of cycles), the scan-rate and the potential window. The thickness (d) of the polymer film can be
roughly estimated from the charge (Q) passed during electropolymerization, according to
Faradays law [243]. By applying the following relation:
d = QM/FAρ
where M is the molar mass of the monomer, F is the Faraday constant (96485 C/m2), ρ is the
density of the polymer, taken to be 1 g/cm3 for polyphenols and A is the electrode surface area.
Herein, the thickness of the polymeric film was about ~2.5 nm. One of the limitations of using
non-conducting polymers, such as phenol and phenol derivatives, is because of their self-limited
growth on the electrode surface and, therefore, resulting in very thin films (< 100 nm) [297].
Here, the target molecule to be imprinted is quite small and, consequently, there is no need to
assemble thicker films. Moreover, previous electrochemical studies have used imprinted
electrodes, modified with non-conducting films, holding ultrathin polymeric layers in order to
improve the sensitivity of the device [298].
Nowadays, impedance techniques are being widely applied to monitor variations in electrical
properties as a direct response of the bio-recognition events that take place at the surface of the
electrodes. Moreover, EIS holds the advantage of not causing any damage to the analyzed
surface and do not introduce any disturbance in the studied system. Herein, along the
fabrication of MIP and NIP, the electrodes were stepwise characterized by means of EIS, using
as redox probe a 5 mM solution of K3[Fe(CN)6] and K4[Fe(CN)6] prepared in 0.1 M phosphate
buffer. Usually, the impedance spectrum includes a semi-circular portion at higher frequencies
and a linear portion at lower frequencies which corresponds to Rct and the diffusion process,
respectively. Figure 6.3 also displays the Nyquist plots showing each step of the modification of
the fabricated NIP (Figure 6.3C) and MIP (Figure 6.3D) electrodes. As expected, the
electropolymerization of phenol resulted in a substantial increase of the Rct value due to the
blocking of electron transfer by the polymeric matrix. Moreover, the results obtained for NIP and
MIP after electropolymerization suggested that the non-imprinted coated electrode exhibited
stronger insulating properties, which constitutes an evidence of the presence of the template
molecule on the imprinted material, which partially hinders the growth of polyphenol film.
Afterwards, the removal step caused a small decrease of the Rct value only on the MIP
modified electrode, suggesting that the formation of the imprinted cavities occurred and,
consequently, electronic diffusion was facilitated. In the literature, there are different approaches
to extract a template molecule from the polymeric matrix of MIP-based sensing devices,
including washing with sulphuric acid, acetic acid, methanol, ethanol, etc. Furthermore, previous
works have already been well succeeded in the removal of template protein by using mixed
organic solvents, such as, PBS:methanol (10:1, v/v), without causing degradation of the
Chapter 6
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polyphenol film [49] or even incubation in ethanol:water (5:1, v/v) mixture [52]. In the current
study, a mixture methanol:water (1:10, v/v) was chosen as the mix solvent with more
satisfactory results, because it was able to efficiently extract the 3-NT template without causing
any degradation or modification to the polymeric material.
6.3.3 Optimization of experimental conditions during MIP assembly
6.3.3.1 Effect of scan-rate and number of electropolymerization cycles
Beside thickness, the density of polyphenol films can also be adjusted by controlling
some variables during electropolymerization, such as, the rate of film growth and the number of
voltammetric sweeps. Hence, comparative studies were performed by using EIS to characterize
the charge transfer properties of different modified electrodes, at various experimental
conditions.
Herein, the MIP-based electrodes were assembled with three different scan-rate values, as
shown in the Nyquist plots of Figures 6.4A-C. Firstly, the only MIP-coated material that
presented a decrease of the Rct value as the result of the creation of imprinted cavities was the
one assembled at a scan-rate of 50 mV/s, while the others suffered an increasing variation.
Moreover, the electropolymerization performed at 50 mV/s scan-rate was responsible for the
formation of a less insulating surface (366 Ω), which could be an advantage for the successful
extraction of the template molecule from the MIP matrix [49]. In addition, Figures 6.4D-F,
illustrate the Nyquist diagrams obtained during the fabrication of the MIP electrodes using 2, 5
and 10 CV cycles, respectively. Generally, the detection capability of a sensor is highly affected
by the number of scanning cycles used in the formation of the polymeric material. For instance,
a lower number of cycles is often associated to a favourable analytical performance whereas a
higher number of voltammetric cycles can led to the formation of a thicker film with less
accessible imprinted sites and, consequently, less sensitivity [102][299]. Furthermore, a high
number of scanning cycles increased the possibility of 3-NT template molecules becoming
trapped in the polymer matrix. Herein, our data showed that the higher % of template removal
was obtained for the MIP prepared with 5 scanning cycles (~10 %), when compared with 2
cycles (~4 %) and 10 cycles (~3 %). Therefore, 5 cycles with a scan-rate of 50 mV/s were
selected for the growth of the non-conductive polymer layer.
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Figure 6.4: A) EIS measurements obtained before and after template removal, at different scan-rates: A)
15 mV/s; B) 50 mV/s and C) 150 mV/s and, with different number of cycles: D) 2; E) 5 and F) 10, recorded
during phenol electropolymerization.
6.3.3.2 Effect of monomer concentration
Among others already mentioned, polyphenol films hold a great advantage as a
biomimetic matrix because it can be easily synthetized by electropolymerization from aqueous
solutions of the monomer. In addition, the thickness of the imprinted layer can also be adjusted
by optimization of the monomer concentration. Herein, EIS was used to follow the construction
of NIP and MIP sensors, with two distinct concentrations of phenol: 1.0 mM (Figure 6.5A) and
0.25 mM (Figure 6.5B). As expected, the Rct values of both NIP and MIP are shifted towards
higher values for increasing phenol bulk concentration. It was interesting to observe that the
stronger insulating character of the non-imprinted coated electrode compared to the imprinted-
device was maintained in both concentrations. The obtained data also have showed that in the
case of 1 mM of phenol (Figure 6.5A) the films formed were quite resistive (4 kΩ < Rct < 7 kΩ),
which may be a limitation in terms of sensor sensitivity.
In addition, the rebinding of 3-NT molecule was quantified by means of differential pulse
voltammetry (DPV). In practice, the current data accounted the occupation of the imprinted
cavities in the MIP electrode by the template molecules, followed by their electro-oxidation.
Figure 6.5C presents DPV calibration curves plotted with the logarithm of current of MIP
sensors against the logarithm concentration of 3-NT. The comparison of the two curves have
showed that a lower concentration of phenol (0.25 mM) enabled lower detection limits when
compared to 1 mM concentration of phenol, which is in agreement with our previous
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conclusions that thicker films hold less number of imprinted cavities and, consequently, less
sensitive responses. So, a concentration of phenol monomer of 0.25 mM was chosen for further
studies.
Figure 6.5: EIS obtained for NIP and MIP sensors at two different phenol concentrations, A) 1 mM and B)
0.25 mM. C) 3-Nitrotyrosine response for both MIP electrodes, obtained from DPV measurements.
6.3.3.3 Effect of imprinted 3-NT concentration
The concentration of the template was another crucial parameter that could influence the
MIP performance. Specifically, the ratio template-monomer of MIP determines the number of
binding sites available for the selective rebinding of the molecules and, consequently, the
detection limit of the biosensor. In parallel, the interaction of these molecules with the phenol
units during the assembly of the polymeric matrix also dictates the formation of stable and
specific binding cavities. Phenol-based MIP films have been widely applied as recognition
matrices in biosensors due to their ability to interact with other molecules through the existence
of the π donor-acceptor interactions [51].
Herein, the effect of 3 different concentrations of 3-NT was investigated during phenol
electropolymerization. Figure 6.6A illustrates the charge variation that MIP electrodes suffer
during the 5 scanning cycles and Figure 6.6B displays the corresponding Nyquist diagrams
obtained after the electropolymerization step. Data showed that for increasing concentrations of
3-NT the tendency was to obtain a less insulating polymeric layer, which is in agreement with
our previous results, where the presence of the template molecule can partially hinder the
growth of polyphenol film. Afterwards, the rebinding of 3-NT molecule was also investigated by
means of DPV for the 3 different concentrations, 0.05 mM (Figure 6.6C), 0.25 mM (Figure 6.6D)
and 0.50 mM (Figure 6.6E). Overall, the higher concentration of imprinted molecule (0.50 mM)
enabled the detection of lower concentrations in comparison with the other imprinted
concentrations of 3-NT, which constitutes a strong indication that using higher concentration of
template molecule results in the creation of more imprinting sites available for the rebinding (see
Figure 6.6F).
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Figure 6.6: A) Charge variation during phenol electropolymerization (5 cycles) obtained from MIPs with
different concentrations of template molecule; B) EIS obtained for the MIPS with different concentrations of
template molecule; DPV measurements after contact with different concentrations of 3-NT for MIPs with C)
0.05 mM, D) 0.25 mM and E) 0.50 mM concentration of template molecule; F) Scheme related to the
distribution of imprinting sites.
Although not presented here, along this optimization study the experimental conditions of the
DPV measurements where finely tuned in order to get the higher response current. Specifically,
one of the great advantages of our approach was the elimination of long incubation periods, by
applying a pre-accumulation potential before each electrical measurement. In all experiments,
the accumulation potential value used before the DPV measurement was 1 V and different
accumulation times were tested, namely 30, 60 and 180 s. Our best electrochemical response
was obtained for 60 s of accumulation at 1 V and so, these conditions were applied along this
work.
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6.3.4 Characterization of the modified paper-electrodes
Raman spectroscopy has been widely used as a popular tool for the characterization and
structural organization of carbon materials. Herein, Figure 6.7A presented the Raman spectra
for the different modifications performed on the paper-based electrodes. Firstly, all samples
presented two well distinct peaks occurring typically in graphite-based substrates, the so-called
G and D peaks, lying at around 1570 and 1310 cm1, respectively. In the clean carbon-coated
electrode, the G band appeared at 1584 cm-1
and was assigned to the C/C stretching in
graphitic materials, composed of sp2 bonded carbon in planar sheets [300]; the D band
appeared around 1307 cm-1
and is often referred to as the disorder band or the defect band. In
order to characterize the level of disorder within the carbon material, the intensity peak ratio
between the D and G bands (ID/IG) was analysed.
Figure 6.7: A) Raman spectra of clean carbon-based electrode, NIP and MIP-modified surfaces. SEM
images of B) NIP and C) MIP materials.
Table 6.2: Analytical data obtained from the Raman spectra related to the different modifications.
Table 6.2 displays the peak positions of the G and D bands in the materials that were
investigated, as well as the corresponding intensities. Although the ID/IG ratios showed a similar
tendency among the different electrodes, it was interesting to observe that the introduction of
Sample D Position
(cm-1
) Intensity
G Position
(cm-1
) Intensity ID/IG
Clean chip 1307.21 347.10 1584.77 269.56 1.29
NIP 1308.13 233.63 1574.36 188.59 1.24
MIP 1310.11 320.74 1572.68 290.53 1.10
Chapter 6
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the polymeric material, on both NIP and MIP cases, resulted in a carbon system with less level
of disorder. Moreover, another typical Raman band of carbon related materials is the 2D band,
usually located around 2700 cm-1
. As can be seen on Figure 7.7A, this band is almost absent in
the spectrum of the clean carbon-coated chip, but on both NIP and MIP materials this signal
appears enhanced, which constitutes another strong evidence of the deposition of the polymeric
material.
Surface morphology of the prepared NIP and MIP electrodes was further studied by means of
SEM, as can be seen in Figure 6.7B and Figure 6.7C, respectively. The main difference
observed between both materials was that the NIP surface presented a higher distribution of
small islets while in the MIP surface the growth of the polymeric film seemed more even and
well spread on the electrode. Moreover, the globular structure obtained for the phenol polymeric
film was in agreement with a previous study, performed in a different substrate material [267].
Although it was not possible to visualize the imprinted cavities, due to the small dimensions of
3-NT molecule, the differences observed in the SEM images can be an indication of the
presence of 3-NT molecule during electropolymerization that resulted in a different structural
polymeric growth.
6.3.5 Performance of the imprinted-sensor
6.3.5.1 Calibration curve
Figure 6.8: Calibration curves corresponding to the response of A) MIP and B) NIP sensors against the
concentration of 3-nitrotyrosine. The inset figure is related to the DPV recordings for each standard
concentration.
The main analytical features of the 3-NT biosensor were evaluated by DPV calibration
curves, carried out under the optimal experimental conditions. Figures 6.8A and 6.8B show the
calibration curves plotted with the logarithm concentration of 3-NT versus the logarithm of
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106
current for MIP and NIP sensors, respectively. As can be seen, a linear tendency was obtained
for both imprinted and non-imprinted materials from 500 nM to 1000 M, but the MIP sensor
showed higher sensitivity (slope = 0.7591), better linearity (r2 = 0.9943) and better
reproducibility (smaller error-bars) in comparison with the NIP. The LOD was 22.3 nM,
calculated as three times the standard deviation from the blank measurement (in the absence of
template molecule). The inset graphic representation illustrated in Figure 6.8A concerns the
DPV data obtained for the MIP-based sensor for the respective concentrations of 3-NT
molecule. The precision of both modified-electrodes, for 3 different independent experiments, is
also presented on Figure 6.8, over the entire concentration range.
Compared with other methods (see Table 6.1), the LOD found here is in agreement with other
works reported in literature involving a molecular imprinting technology coupled with
electrochemical sensing. Interestingly, when reviewing the construction of these imprinted-
based sensors, it was generally found that the use of amplification steps, such as, the
incorporation of nanomaterials, was required to achieve such low detection limits. Besides,
another relevant advantage of our electrochemical device is the use of a label-free approach
that enables decreasing substantially the analysis time and, consequently, becoming more
affordable. The use of a paper support is also a great advantage in terms of cost.
6.3.5.2 Urine samples
Figure 6.9: Calibration curves corresponding to the response of A) MIP and B) NIP sensors against the
concentration of 3-nitrotyrosine, performed in 1:10 diluted human urine samples.
In order to evaluate the applicability of the proposed MIP-modified electrode in real
samples, the electrochemical readings of 3-NT standards were performed in a background of
human urine samples. Thus, diluted blank urine samples (1:10 in phosphate buffer) were spiked
with different concentrations of 3-NT and analysed by means of DPV measurements. As can be
seen in Figure 6.9A, the MIP material enabled the quantification of 3-NT from 5 m to 1 mM,
with RSD ranging 1 to 2 %. In contrast, the NIP material (Figure 6.9B) showed a non-
reproducible response along a narrow range of concentrations, which could be a strong
Chapter 6
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indication of the non-specific adsorption phenomena occurring at the surface of the polymeric
material.
Overall, these results demonstrated the capability of this imprinted-modified sensor assembled
on paper support to sensitively quantify 3-NT in urine samples. Despite the existence of few
sensor devices with similar LODs for the detection of 3-NT, our purpose herein was the
development of a quick, facile, reproducible and low-cost paper-based biosensor suitable to be
applied as a portable tool in POC screening.
6.4 CONCLUSIONS
This study reports a label-free electrochemical biosensor platform for in-situ detection of
3-NT. Herein, the application of an imprinted material could enable specific rebinding of
molecules into the imprinted cavities of MIP-modified films, improving selectivity features of the
biosensor device. The optimization of some experimental parameters during
electropolymerization, such as, scan-rate, number of cycles, monomer and template
concentration, among others, allowed to finely tune the thickness and growth of the polymeric
matrix. In sum, the imprinted-based sensor showed high sensitivity and selectivity towards 3-NT
over a wide range of concentrations and was successfully applied to the analysis of biological
samples. Although the obtained LOD was not the lower value found in the literature, the
electrochemical performance of this paper-based biosensor greatly satisfies the requirements of
a selective, reproducible, disposable and sustainable sensing platform that can be easily and
costly miniaturized in the future.
Chapter 6
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Chapter 7
109
CHAPTER 7
7 Synthesis and characterization of core-shell
magnetic nanoparticles
7.1 INTRODUCTION
As previously mentioned, MNP are being widely used as an efficient solid platform for the
isolation, purification and pre-concentration of target molecules due to their straightforward
application, hindering occasional losses of analyte [110]. Moreover, through the facile
application of an external magnetic field, these nano-sized particles can be easily retained
during the application of a magnetic force and then, quickly release after the exposure, without
suffering significant changes. In addition, magnetic-based delivery approaches have the
intention to associate drugs with magnetic vehicles in order to concentrate and, afterwards,
release the same drug at the local of interest [301]. Thus, the incorporation of magnetic
nanoparticles as novel sensing substrates holds great advantages, such as, the use of small
sample volumes and easy-to-use procedures, which also diminish the analysis time. In addition,
the properties of nanomaterials are usually size dependent and, often, they exhibit unique
physical and chemical properties when compared to the bulk counterparts [302]. Herein, the
integration of magnetic nano-sized structures into the assembly of biosensor devices,
specifically electrochemical-based sensors, can be used as a way to amplify the signal. So,
under this scope, the development of core-shell nanoparticles enables the combination of the
magnetic properties of the core combined with the functionalized properties of the shell's
material.
In parallel, silica coated core–shell structured particles constitute a popular choice, mostly
because silica is nontoxic, biocompatible and can be easily functionalized with different
functional groups [303]. Among the wide range of applications of silica-based nanoparticles, the
most common are construction materials, drug delivery systems, preparation of polymer
composites and also treatment of diseases [304][305][306]. For instance, fluorescent iron oxide-
silica nanoparticles have been synthesized by introducing organic fluorescent dyes into the
silica shell, resulting in a unique combination of magnetic and optical characteristics [224]. In
addition to biocompatibility and hydrophilicity, the formation of a silica shell can also prevents
Chapter 7
110
the oxidation of magnetic particles and, simultaneously, enables different approaches for
surface modification through the covalent attachment of specific ligands on the surface of these
nanostructured materials. Another approach that can be employed to preserve the integrity of
the magnetic core, is the layer-by-layer growth of noble metals such as Ag, Au and Cu [307].
The reason why such metals were used is the fact that they are biocompatible, resistant to
physiological conditions and their shells do not allow agglomeration of the cores. Nevertheless,
the incorporation of noble metals into nano-sized structures is typically limited to expensive and
more complex fabrication technologies. Thus, herein, the silica shell not only prevents the
magnetic cores from oxidation, but also provides a surface with the ability to incorporate
additional materials holding other functionalities, such as, optical or electrochemical behaviour.
There are two different routes that have been commonly explored in order to generate silica
coatings on the surfaces of iron oxide particles. The first method relies on a sol-gel process,
also called the coprecipitation method or the well-known "Stöber process", in which silica is
formed in-situ through the hydrolysis and condensation of a sol-gel precursor [302]. Generally,
the Stöber method was employed to cover nanoparticles that hold the ability to become
disperse in polar solvents, like water or ethanol. In the late 1960s, Stöber et al. developed a
process capable of forming controlled silica particles (500 nm - 2 μm) using tetraethoxysilane
(TEOS) as silica precursor, water, ethanol and ammonia [308]. Briefly, silica particles had a
narrow size distribution and could be adjusted by controlling the pH of solution, composition of
reactants and temperature. One of the main advantages of the sol-gel technique is that it offers
a low temperature method for synthesizing and mixing inorganic with organic materials. The
other method is based on microemulsion synthesis, in which micelles or inverse micelles,
stabilized by surfactant film, are used as nano-reactors to confine and control the formation of
nanomaterials [216].
Figure 7.1: Schematic representation of the different steps related to the synthesis of the core-shell
magnetic nanoparticles.
Chapter 7
111
In this work, silica (SiO2) was chosen as the surface coating of magnetic nanoparticles because
probe molecules can be easily encapsulated into the silica shell, allowing in parallel the
functionalization of the surface for bioconjugation and targeting for biological applications.
Various labelling agents can be incorporated into nanoparticle structures yielding materials with
a broad spectrum of applications, such as, fluorescent beads, redox probes, dye-labeled
nanoparticles or semiconductor quantum dots. For instance, a study have performed the
synthesis of bifunctional nanoparticles, SPIO@SiO2(FITC) with both fluorescent and magnetic
properties, in order to develop contrast agent-loaded nanoparticles with which stem cells could
be efficiently and harmlessly labeled and then imaged with a clinical MRI analyzer at low cell
number [309]. Thus, herein, we describe a sol-gel method based on the hydrolysis of TEOS for
coating iron oxide nanoparticles with an uniform silica shell, enabling the encapsulation of redox
probe species (see Figure 7.1).
Novel strategies for signal amplification have been extensively investigated, specifically, for the
development of ultrasensitive immunoassay. Thus, combining Fe3O4 magnetic nanoparticles
with silicon dioxide shell and gold nanoparticles (Fe3O4@SiO2@Au) enabled the fabrication of
an electrochemical immunoassay for the detection of C-reactive protein in serum samples [310].
Due to their large surface area, the functionalization of magnetic nanoparticles holds a great
potential application in the biotechnology field, allowing the pre-concentration of analyte
molecules. In a similar way, electrochemical redox-active species, such as, methylene blue
(MB) have been incorporated with gold-modified magnetic nanoparticles for signal generation
and amplification [311]. Interestingly, functionalized nano Fe3O4 particles (magnetic cores) have
also been combined with molecularly imprinted polymer to accomplish adsorbent materials for
imidacloprid target with amplified signal [312]. In sum, herein we present the fabrication and
characterization of Fe3O4@SiO2 nanoparticles modified with an redox-active probe with
potential to be used as labels in an amplified multiplexed array.
7.2 EXPERIMENTAL SECTION
7.2.1 Reagents and Materials
All chemicals were of analytical grade and used as supplied without further purification.
PBS (0.01 M, pH 7.4) solutions were prepared with ultrapure water Milli-Q laboratory grade.
NADH (grade II, disodium salt 98%), tetraethyl orthosilicate (TEOS), 3-triethoxysilylpropylamine
(98% purity, APTES), citric acid and ethanol absolute (≥99.9% p.a.) were obtained from Sigma-
Aldrich; hexaammineruthenium (III) chloride (Ru(NH3)6Cl3) was from Acros Organic; iron (II)
sulfate 7-hydrate (FeSO4·7H2O) was from Panreac; iron (III) chloride hexahydrate (FeCl3·6H2O)
was from Scharlau; PBS tablets from Amresco; sodium hydroxide (NaOH) from EKA;
ammonium hydroxide (25% p.a.) from Fluka and sodium docedyl sulfate (SDS) from TCI.
Chapter 7
112
The exact pH values were measured with a pH meter (Crison Instruments, GLP 21 model). All
experiments were carried out at ambient temperature.
7.2.2 Apparatus
The electrochemical measurements were conducted with a potentiostat/galvanostat from
Metrohm Autolab and a PGSTAT302N with a FRA module, through an interface switch box from
BioTid Eletrónica and controlled by ANOVA software. Commercial carbon-screen printed
electrodes (SPEs-DRP-C110) were purchased from DROPSENS. All the presented potential
values are related to the silver pseudo-reference of the SPEs.
In order to follow the introduction of functional groups during each step of modification of the
nanoparticles, FTIR measurements were performed in the ATR mode, using a Thermo Scientific
Smart iTR Nicolet iS10, coupled with sampling accessory of diamond contact crystal, also from
Nicolet.
To assess the morphology of the surface as well as the distribution of sizes of the nanoparticles,
both SEM and TEM techniques were employed. For TEM analysis, visualization was carried out
on a JEOL JEM 1400 TEM at 120kV (Tokyo, Japan). Images were digitally recorded using a
CCD digital camera Orious 1100W Tokyo, Japan at the HEMS / i3S of the University of Porto.
SEM studies were performed on a FE-CryoSEM/EDS, from JEOL JSM 6301F, Oxford INCA
Energy 350, Gatan Alto 2500 microscope, operating at 15 and 25 kV.
The thermal behaviour of the magnetic nanoparticles was assessed by performing
thermogravimetric (TG) experiments with an Hitachi TG/DTA/7200 analyzer.
7.2.3 Synthesis of core-shell nanoparticles
7.2.3.1 Fabrication of iron oxide nanoparticles
The preparation of MNPs was performed by a chemical co-precipitation process involving
Fe2+
and Fe3+
ions (1:2 molar ratio) in an alkali medium (see Figure 7.2). Briefly, 20 ml of 0.55 M
Fe3+
and 0.275 M Fe2+
solution was prepared with ultrapure water in an Erlenmeyer flask.
Afterwards, this mixture was kept with magnetic stirring at 40ºC under nitrogen atmosphere until
complete dissolution of the reagents. An inert atmosphere (without oxygen) is essential to
obtain pure magnetite particles. Then, sodium hydroxide (25 wt%) was added to the reaction
mixture, leading to the formation of a green rust at the early stage of precipitation, followed by a
black colour after precipitation process was completed (~30 min). The surface of the MNP was
stabilized with citric acid (2 M) and the reaction was maintained at 40 ºC for 90 min with
continuous stirring. Finally, the reaction mixture was left cooling and afterwards, the
nanoparticles were washed several times with ultrapure water until a pH of 7 was obtained. The
wet suspension was stored in the fridge.
Chapter 7
113
Figure 7.2: Synthesis reaction of the iron-oxide nanoparticles via co-precipitation method.
7.2.3.2 Preparation of iron oxide-silica core-shell
SiO2-MNPs were prepared following the conventional Stöber method, as described in the
literature with some modifications. Firstly, 125 mg of the as-prepared iron oxide solution was re-
suspended in 20 mL of water with continuous magnetic stirring and then, 0.1 % (wt) solution of
SDS was added. Afterwards, the stirring process was continued followed by the addition of 10
mL of ammonia solution. In parallel, the redox probe (NADH or Ru(NH3)6Cl3) was dispersed in a
mixture water:ethanol and then, a small volume of TEOS was added. Finally, the previous
mixture was added to the previous nanoparticles solution and magnetic stirring was continued
for more 3 hours, at room temperature. Different concentrations of TEOS precursor were
investigated. The modified nanoparticles were re-suspended and washed several times with
ultrapure water and ethanol.
In order to prepare amino-functionalized MNPs, 100 mg of SiO2-MNPs were firstly dispersed in
a mixture 1:1 (v/v) ethanol: water with continuous magnetic stirring at room temperature.
Afterwards, APTES (2 %, v/v in ethanol) was introduced into the solution and left stirring for 5 h
at room temperature. In the end, the modified nanoparticles were removed from the mixture and
thoroughly washed with ultrapure water and ethanol.
7.2.4 Characterization of the modified nanoparticles
Each step along the synthesis and modification of nanoparticles was followed by FTIR,
TEM, SEM and TG analysis. Regarding FTIR experiments, spectra were collected under room
temperature, after background correction. The number of scans for sample was set to 32 and
the resolution was fixed at 8. FTIR data analysis was performed with OMNIC software. For TEM
analysis, the nanoparticles were dispersed in ultrapure water, sonicated for 10 min and
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afterwards, 10 µL of samples were mounted on Formvar/carbon film-coated mesh nickel grids
(Electron Microscopy Sciences, Hatfield, PA, USA) and left standing for 2 min. The liquid in
excess was removed with filter paper. Before SEM analysis, nanoparticle samples were dried at
80 ºC, for 24 hours. Finally, TG experiments were performed in aluminium crucibles with a
heating rate of 5ºC/min, from +30 ºC until 550 ºC, under a nitrogen atmosphere of 200 mL/min.
7.2.5 Electrochemical assays
Electrochemical measurements for characterization of the modified magnetic
nanoparticles were performed by using different electrochemical techniques, such as, CV and
SWV. For CV assays, the potential was scanned from 0.7 to +0.7 V, at 50 mVs-1
, in a magnetic
nanoparticle solution prepared in 0.01 M PBS solution, pH 7.4. The measurements were
performed by covering the three electrodes with 80 l of the appropriate solution.
Initially, NADH and Ru(NH3)6Cl3 were selected as electrochemical probes and their redox
behaviour was investigated by SWV, over a wide potential range. Afterwards, for the modified
M-NP containing labeled Ru(NH3)6Cl3, SWV was performed from -0.40 V to +0.20 V, at a very
low scan-rate. For this procedure, the nanoparticles were prepared in PBS solution, transferred
to the SPEs and then concentrated, with the aid of a magnet under the sensing platform. The
magnet is maintained during the entire electrochemical reading.
7.3 DISCUSSION AND RESULTS
7.3.1 Synthesis of iron oxide nanoparticles
Herein, we have performed a chemical co-precipitation of Fe2+
and Fe3+
under alkaline
media for the preparation of nanosized Fe3O4 particles. In order to assess the size distribution
and the morphology of the nanoparticle, TEM analysis was performed. As shown in Figures
7.3A and 7.3B, MNPs were about 20 nm in average size presenting a nearly-spherical
morphology with a quite narrow size distribution. Although not presented here, the incorporation
of citric acid during the synthesis of the magnetic particles was crucial to stabilize the surface of
MNPs in an aqueous dispersion, hindering the formation of agglomerates. This outcome is in
accordance with literature whereas agglomeration is a common result of the synthesis by the
co-precipitation technique if a capping agent, like citric acid, is not used [313][314].
Furthermore, the resulting black precipitate (Fe3O4) obtained presented good magnetic
properties in the presence of a magneto, as displayed in Figure 7.3C.
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Figure 7.3: (A-B) TEM images of iron-oxide nanoparticles at different magnifications and C) image of the
MNPs under the application of a magnetic field.
7.3.2 Fabrication of iron oxide-silica core-shell
Along this study, two different redox-active species were tested that includes, ruthenium
(Ru3+
/Ru2+
) and NADH (NAD+/NADH). In order to assess the electrochemical behaviour of these
two probes, SWV was the chosen technique mostly due to their high sensitivity and quick
analysis time. Figure 7.4 shows the voltammograms for both individual solutions prepared in
PBS at a pH of 7.4. The oxidation peak of ruthenium was around -0.25 V while NADH oxidation
occurred at +0.50 V. These two probes were selected because their oxidation peaks were well
separated and thus, no overlapping would take place.
Figure 7.4: Square-wave voltammograms concerning the two (individual) redox probes ruthenium and
NADH, in PBS at a pH 7.4, applied in a clean, bare carbon-SPE.
Chapter 7
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Among others, magnetic Fe3O4 nanoparticles hold important characteristics, such as,
superparamagnetic property, good biocompatibility, low toxicity and facile preparation, making it
a popular choice as anchoring platforms. Under this scope, coating this nanostructured
materials with redox-active species encapsulated in inert shells is an interesting approach to
preserve the functional characteristics of the species. Moreover, coating can also prevent
nanoparticles from contacting other nanoparticles, so that aggregation of nanoparticles, which
spoils blood flow, can be hindered [315]. Herein, SiO2-MNP were synthesized based on a sol-
gel chemistry in order to incorporate ruthenium or NADH probe in the silica shell. Similarly, a
variety of fluorophore compounds can be encapsulated within the silica coating, resulting in a
fluorescence emission in the near-infrared region that is beneficial for the analysis of biological
tissues [316]. Another strategy used to synthesize labeled-nanoparticles for biosensing
purposes is the incorporation of electroactive metals into phosphate-based nanoparticles to be
used in an electrochemical multiplex biosensor [317]. Moreover, the fabrication of different
electrochemical redox-active species based on HAuCl4 and Na2PdCl4 as co-oxidating agents
and aniline derivatives as monomers have also been reported [318].
The sol-gel-based Stöber synthesis is usually described with three reaction as shown below for
silicon: hydrolysis, alcohol condensation and water condensation. The reaction mechanism is
well described in the literature [308]:
1. Hydrolysis
Si(OCH2CH3)4 + H2O (CH3CH2O)3Si-OH + CH3CH2OH
2. Alcohol condensation
Si(OCH2CH3)4 + (CH3CH2O)3Si-OH (CH3CH2O)3Si-O-Si(OCH2CH3)3 + CH3CH2OH
3. Water condensation
(CH3CH2O)3Si-OH + (CH3CH2O)3Si-OH H2O + (CH3CH2O)Si-O-Si(OCH2CH3)3
Finally, a SiO2 network is produced:
Si(OCH2CH3)4 + H2O SiO2 + 4CH3CH2OH
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Figure 7.5: SEM images of the MNPs A) non-modified, modified with B) SiO2 with NADH and C) SiO2 with
ruthenium; EDS spectra of the MNPs D) non-modified, modified with E) SiO2 with NADH and F) SiO2 with
ruthenium.
To evaluate the morphology and the size distribution of the as-prepared nanoparticles, SEM
was performed. Figure 7.5A, 7.5B and 7.5C displayed the surface of the nanoparticles and the
size measurements concerning magnetic nanoparticles non-modified, SiO2-M-NP with NADH
and SiO2-M-NP with ruthenium, respectively. As expected, a substantial increase of the
diameter was observed which constitutes a good indication that the silica coating was
successfully incorporated into the particle. Moreover, it seems that the spherical shape of the
modified particles was also maintained. Although the size distribution in the modified magnetic
nanoparticles was not very narrow, it was interesting to observe that the type of redox probe
used during the formation of the silica shell had a great impact on the size of the obtained
nanoparticle. In addition, the presence of silicon element in the silica-modified magnetic
nanoparticles was also confirmed by Energy Dispersive X-Ray Spectroscopy (EDS) analysis, as
can be seen in Figures 7.5E and 7.5F, in contrast with the absence of this element in Figure
7.5D (non modified MNPs). Moreover, EDS results also indicated a decreasing in the silicon
content when the ruthenium probe was incorporated, thereby confirming the presence of the
ruthenium element.
Covalent immobilization of simultaneous electroactive molecules into modified Stöber silica
nanoparticles have been combined in a unique environment for biotechnology purposes.
Concerning their catalytic properties, previous electrochemical studies have reported that the
redox potentials of the ferrocene and ruthenium(II) complex species immobilized on the silica
particles are not very different from those of the corresponding solution species [319]. Figure
7.6A shows the application of CV to study the electrochemical behaviour of the magnetic
nanoparticles. For this study, each type of nanoparticles was dispersed in PBS pH 7.4 and a
drop of this solution was used to cover the 3-electrodes of carbon-SPEs. By looking to the
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118
voltammograms, we can verified that non-modified magnetic particles do not display any
electrochemical behaviour, while the nanoparticles modified with the redox ruthenium showed
several oxidation and reduction peaks. Specifically, comparing this data with the solution specie
presented previously, the characteristic oxidation peak around -0.25 V was maintained.
Moreover, these results also showed that the silica-modified nanoparticles prepared in the
presence of NADH did not exhibit any redox peak.
Figure 7.6: A) Cyclic voltammogram applied in PBS at pH 7.4 of the different types of MNPs; B) images of
the wet suspension of the MNPs B1) non-modified, modified with B2) SiO2 with NADH and B3) SiO2 with
ruthenium and TEM images of the MNPs C1) non-modified, modified with C2) SiO2 with NADH and C3)
SiO2 with ruthenium.
Figure 7.6B displays the magnetic nanoparticles in aqueous solution and we observed that, in
contrast with the bare magnetic particles, the modified nanoparticles had the tendency to
deposit in the bottom of the flask, which can be the result of more heavy particles. Although
there are some evidences that magnetic properties of coated materials can decrease
substantially upon the increase of outer shell thickness [224], herein, the magnetic properties of
the obtained nanoparticles were preserved.
TEM analysis (see Figure 7.6C) pointed out a great level of agglomeration of the magnetic
nanoparticles after the silica sol-gel step, which constitutes a common issue during this type of
synthesis [302]. Despite the aggregation, Figures 7.6C2 and 7.6C3 clearly showed a darker core
correspondent to magnetic nanoparticles surrounded by a large silica shell. So, although the
silica modification was successfully achieve, more optimization studies are needed in order to
tune the coating growth, in order to prevent agglomeration problems. Moreover, among other
factors, the thickness of silica coating can be easily controlled in the range of 2-100 nm by
changing the concentration of the TEOS precursor [224]. In sum, for the following optimization
studies, silica-modified magnetic nanoparticles with ruthenium probe was the chosen system
due to the fact that enabled an electrochemical signal.
Chapter 7
119
7.3.2.1 Optimization of the sol-gel process
Solvent and surfactant effect:
More systematic studies have been performed in order to understand how polydispersity
of nanoparticles in a synthesis process can be highly influenced by the synthesis kinetics [320].
So, they have concluded that a faster reaction will result in a higher polydispersity if allowed to
react for longer periods. An overview of the literature have showed that the key factors
responsible for controlling the particle size and distribution during silica sol-gel synthesis include
ammonia, solvent (water versus alcohol), concentration of TEOS and temperature. Herein,
ammonia was used as the catalyst of the hydrolysis and condensation of the silicate species
and apparently, has great influence in the morphology of the particles [224][321].
Figure 7.7: TEM (A) and SEM (B) imaging of the silica-based nanoparticles prepared (1) without ethanol
and SDS, (2) without ethanol and with SDS and (3) with ethanol and SDS.
In this work, the effect of using a surfactant (SDS) and also the influence of water and/ or
ethanol presence was investigated through the analysis of TEM (see Figure 7.7A) and SEM
(see Figure 7.7B). It is well known that the use of surfactants with relatively high concentrations
is usually necessary to prevent aggregation of nanoparticles. So, it was observed that the use of
a small concentration of SDS during the sol-gel synthesis (see Figure 7.7A2) resulted in less
particle aggregation with a more homogenous distribution of the magnetic cores in comparison
with the non-surfactant system (see Figure 7.7A1). Moreover, another important parameter with
great impact in the size distribution of silica-based nanoparticles that was also investigated
along this study is the effect of ethanol. An important role of ethanol presence in the medium is
the enhancement of TEOS solubility during the polymerization reaction. Thus, both Figures
7.7A3 and 7.7B3 displayed nanoparticles dispersions with less agglomeration phenomenon and
enabling well-defined particles. In sum, the optimal nanoparticle system was obtained when
using ethanol as solvent and small amounts of SDS surfactant.
Chapter 7
120
TEOS concentration:
In this work, we investigated the effect of the concentration of TEOS precursor not only
on the size distribution of the nanoparticles but only on their electrochemical performance. The
TEM analysis presented in Figures 7.8D, 7.8E and 7.8F showed that the particle size increased
with increasing concentration of TEOS, which is in agreement with similar works [322].
Moreover, in our case, more important than the size of the nanoparticle is the growth of the
silica-based coating and, consequently, the level of aggregation between those particles. So,
the analysis of SEM data indicated that a better dispersion was obtained for the sol-gel
synthesis conducted at lower TEOS concentration, enabling the formation SiO2-MNP with a
regular spherical shape and diameters around 20 nm. Nanoparticle solutions prepared with the
same concentration were prepared in PBS and analysed by SWV. As can be seen in Figures
7.8A-C, the higher amplitude of the oxidation peak was accomplished for the lower
concentration of TEOS, which can be a strong indication that the use of more TEOS resulted in
a thicker silica coating and thus, hindering the electrochemical signalling.
Figure 7.8: Square wave voltammograms of the silica-based MNPs synthesized with increasing
concentration of TEOS A) 0.1 mL, B) 0.5 mL and C) 1.0 mL; TEM images of the silica-based MNPs
obtained with different TEOS concentration D) 0.1 mL, E) 0.5 mL and F) 1.0 mL.
Chapter 7
121
7.3.2.2 Functionalization with APTES
As mentioned before, besides behaving as a biocompatible and protective shell, the
formation of SiO2 coatings offers many possibilities for surface modification through the covalent
attachment of specific functional groups. Moreover, the application of magnetic nanoparticles
has shown to be an effective approach for immobilization of different redox probes on the
nanoparticle surface in order to generate an electrochemical signal. Along this study,
optimization of the experimental conditions during the nanoparticles synthesis and the choice of
solvent were crucial to achieve modified nanostructured that were chemically stable, well-
disperse in liquid media and uniform in both size and shape. Figure 7.9 shows the multi-step
procedure for the fabrication of magnetic nanoparticles and their application as label-less
nanoprobes in electrochemical platforms for biosensing purposes. For instance, a sandwich-
type electrochemical immunosensor combined with gold magnetic nanoparticles have been
designed for simultaneous detection of four different antigen species [323].
Figure 7.9: Schematic illustration of the fabrication procedure of the core-shell magnetic nanoparticles and
their application as electrochemical probes.
After selecting the optimum condition of the precursor composition, and its effect on the SiO2NP
size distribution, APTES was used as a chemical modifier to form amino-end groups onto the
surfaces of the silica coated nanoparticles. This approach is commonly used to facilitate the
combination of these nanostructured materials with biomolecules, such as, drugs, proteins or
antibodies [324]. Specifically, IgG antigens can be covalently attached via their carboxyl group
activated by EDAC/NHS chemistry to the amino groups of silica coated MNPs. Figures 7.10A
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122
and 7.10B shows the SWV measurements related to the silica coated magnetic nanoparticles,
in order to evaluated their electrochemical stability. Firstly, it can be seen that the three
consecutive measurements for each step displayed a quite reproducible behaviour. In addition,
the application of a magnetic field enabled an enhancement of the electrochemical signal as a
result of the attraction an concentration of the nanoparticles in the transducer surface. As can
be seen in Figure 7.10C, after the chemical modification of silica surface with APTES, the peak
amplitude obtained during the SWV measurement decreased substantially which constitutes a
strong indication that the amino groups were successfully introduced in the nanoparticles
surface.
Figure 7.10: Square wave voltammograms of the developed silica-based MNPs at different stages of
fabrication: A) TEOS with ruthenium, B) effect of a magnetic field on the previous MNPs and C) the effect
of functionalization with APTES; D) Image of the electrochemical measurement of MNPs using a magneto.
Generally, the chemical modification of silica surface using APTES agents can be performed via
aqueous or non-aqueous system. The non-aqueous system is the most common used because
silanes, like APTES, carries amine groups that can undergo uncontrollable hydrolysis and
polycondensation reactions in the aqueous conditions [302]. Herein, the use of ethanol solvent
during the coupling reaction with APTES was meant to achieve a better control of the reaction
conditions. Although there are some studies where the formation of the APTES coating could
hindered the magnetic properties of the nanoparticles [325], in our work the APTES-modified
nanoparticles exhibited a strong magnetization in the presence of a magnetic field. As shown in
Figure 7.10D, they displayed a good magnetic response being easily attracted by a magnet.
Chapter 7
123
7.3.3 Characterization of the modified-nanoparticles
During each step of the synthesis and modification of the iron oxide core-shell
nanoparticles, FTIR and TG analysis were performed in order to confirm each modification. The
FTIR spectra of bare MNPs, Fe3O4@SiO2 MNPs and Fe3O4@SiO2@APTES MNPs are
presented in Figure 7.11. The characteristic band attributed to the Fe-O stretching vibration
mode of the magnetic nanoparticles appeared around 550 cm-1
, which is in agreement with
previous works [224]. An intense band was also observed at 3300 cm-1
that may be assigned to
the existence of non-dissociated OH groups from citric acid capping. In addition, the peak at
1650 cm–1
may be assigned to the symmetric stretching of OH from COOH group, as a result of
the binding of citric acid radical to the iron-oxide nanoparticles. Afterwards, the silica network
was adsorbed on the magnetite surface through the formation of Fe–O–Si bonds. Thus, the
adsorption of silane polymer onto the surface of the magnetic nanoparticles was confirmed by
the strong peaks near 1100 and 900 cm−1
assigned to Si–O–Si and Si–O stretching vibrations,
respectively. This outcome is in accordance with other studies that confirms the condensation
reaction between hydroxyl groups on the magnetite nanoparticle surface and the alkoxysilane
molecule [326]. The introduction of APTES can be confirmed by the broad FTIR absorption
band around 3300 cm-1
from N–H stretching vibration and also a sharp peak was observed at
1600 cm-1
assigned to the NH2 bending mode of free NH2 group [325][327].
Figure 7.11: FTIR spectra of the magnetic nanoparticles at each step of the fabrication and
modification reaction: blue line (MNPs); green line (Fe3O4@SiO2 MNPs) and red line
(Fe3O4@SiO2@APTES MNPs).
Chapter 7
124
Figure 7.12 shows the TG analysis weight loss curves correspondent to the iron oxide magnetic
nanoparticles during each modification step. Firstly, as can be seen, Fe3O4@SiO2 and APTES-
coated Fe3O4@SiO2 MNPs presented a distinct mass-loss profile compared to that obtained in
the naked Fe3O4 core. The initial weight loss from 80-180 ºC is generally caused by the removal
of the adsorbed (free) water and ethanol molecules, which is in agreement with similar works
[230]. Afterwards, an accentuated decrease of mass was observed for the bare magnetic
nanoparticles in the range 180-300 ºC that can be attributed to the decomposition of the citric
acid capping [328]. In addition, the sharp increase in weight loss that occurred in the region
180-400 ºC is the result of the thermal decomposition of organic materials that composed the
shell of the magnetic nanoparticles. It was observed that no significant differences were
detected for both Fe3O4@SiO2 and APTES-coated Fe3O4@SiO2 MNPs, being the total weight
loss amounted to 11.7% and 11.2%, respectively. This outcome can be explained due to the
very small amount of APTES on the surface of the nanoparticles, as observed in other similar
studies [326]. After the organic shell had been completely decomposed, the residual mass was
mainly Fe3O4.
Figure 7.12: TG analysis of the different stages of the magnetic nanoparticles.
7.4 CONCLUSIONS
Firstly, this work have presented a promising approach to synthesize and modify
magnetic nanoparticles to be incorporated in multiplexed arrays as electrochemical nanoprobes.
The effect of the redox-active specie, the concentration of the organosilane precursor and the
use of a surfactant were important parameters that affected the size and electrochemical
performance of the nanosized particles. Moreover, the growth of an inorganic structure, such
Chapter 7
125
as, silica, was successfully accomplished enabling the encapsulation of ruthenium in the shell,
while preserving its electroactive behaviour. Preliminary results have also showed that the use
of magnetic nanoparticles as a sensing platform allowed the amplification of signal. In the
future, this dual-stimulus nanoparticles are proposed to be included in the design of a novel
multiplexed array sensor towards various OS biomarkers. To accomplish this goal, another
redox probe, different from ruthenium, will be encapsulated into the SiO2-MNP core-shell
structures in order to get two different electrochemical nano-labels for the simultaneous
detection of MDA and 4-HNE biomarkers.
Chapter 7
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Chapter 8
127
CHAPTER 8
8 Conclusions and future perspectives
Overall, the design of the electrochemical sensors developed along this thesis have
showed to accomplish the same requirements of the current methodologies used, mainly,
sensitivity, reproducibility and selectivity, combined with the advantages of using environmental
friendly materials, with a cost-effective analysis suitable for POC screening. Although some
limitations were found during the work associated to the function of some devices, the
introduction of nanomaterials and also the incorporation of biomimetic films enabled obtaining
the needed standards of performance required for these novel electrode systems.
In the future, it is important to keep in mind that sensitivity is also controlled by the type and
nature of the transducer. Thus, herein the electrocatalytic behaviour of the selected material can
also influence the LOD of the sensor. Specifically, the comparison of all sensing devices
developed along this thesis showed that, as expected, the lowest LOD value was accomplished
by the gold electrode (chapter 3) against the other carbon materials (chapters 4-6). Despite this
great advantage, the use of this kind of gold electrodes imposes multi-step and time-consuming
cleaning protocols that are essential to assure the reproducibility of the experiments. Therefore,
from the point-of-view of biosensor application in routinely conditions, rapid and practical
devices are highly preferred. Moreover, it is important to understand that this response is not
directly transposed into commercial screen-printed electrodes of gold, as these are made by an
composite of gold and not the pure metal.
One of the most important issues that influenced the electrochemical performance of the MIP-
based sensing materials was the assembly of very thin films due to the need that ideally the
recognition event should take place as closer as possible to the electronic transduction surface.
So, regarding the synthesis strategy of the biomimetic polymer, electropolymerization was
acknowledged as the most suitable technique since it enabled a high adherence to the
transducer substrate, an easy control of the film thickness and growth, with the possibility to
work at mild conditions, such as, aqueous solutions and ambient temperature. In addition,
phenol was identified as a suitable material to produce the polymeric matrix that creates the
recognition sites, due to its high stability and specific affinity to the target molecules.
Along the construction, characterization and application of each of the electrochemical sensors
developed herein for the detection of different OS biomarkers, important complementary
Chapter 8
128
techniques, such as, Raman spectroscopy, SEM and FTIR were employed in order to confirm
the formation and growth of the different materials. In addition, an innovative strategy based on
FITC-labelled 8-OHdG antibodies was used to verify the existence and distribution of the MIP
cavities along the surface by specific electron microscopy technique.
In parallel, an intensive voltammetric investigation of the paper-based printed-electrodes has
confirmed their suitable and reliable electrochemical performance. The simple procedure of
microfabrication of this transducer platform started with the hydrophobization of paper, followed
by the application of a suitable conductive ink into the planar substrate materials and then, a
proper thermal cure. Once again, this quick and facile procedure opens new opportunities to
create and implement routine and low-cost assays that can be used in detection, monitoring and
prevention context. Herein, the first attempt of homemade carbon-based electrodes constructed
onto paper substrates (chapter 4) was accomplished with a system of three electrodes,
separately. Although the developed sensor hold important features, such as, good stability and
reproducibility, the capability of regeneration and quick analysis time, the evolution for an
integrated 3-electrode system applied in the same platform was necessary (chapter 5 and 6).
Overall, the combination of a molecular imprinting strategy responsible for tuning the specific
binding of the target molecule with the fabrication of a flexible paper-based sensor seems to be
the best outcome of this work. Thus, the biosensor described in chapter 6 enabled the required
selectivity and sensitivity and, at the same time, the reproducibility issues of the commercially
available SPEs was overcome, with the introduction of a substrate that was environmental-
friendly. Furthermore, another great advantage of these electrodes is also their ability to be
drawn in different shapes and sizes, using different kinds of materials. In this way, for enhanced
results, the optimization of the dimensions used along this work for the paper integrated with 3-
electrode system could be tried out in order to obtain a better analytical response.
The fabrication and functionalization of iron-oxide core-shell nanoparticles has proven to be a
potential approach to be incorporated in novel signal amplified multiplexed array. Ideally, two
magnetic nanoparticles may be functionalized with different labels and targeted to different
biomarkers and, afterwards, applied together in the sensing surface for simultaneous
electrochemical detection. Herein, silica coating was successfully used not only to encapsulate
an electrochemical redox-active element but also allowed to stabilize the magnetic
nanoparticles in solution. It constitutes a promising way to fabricate novel sensing platforms
holding dual-stimulus that are able to eliminate, or hinder, some usual artefacts present during
the samples analysis.
Looking towards the future, additionally to ruthenium-Fe3O4@SiO2@APTES MNPs, a different
redox specie can be also used in order to get different electrochemical labels that can be
applied in a multiplexed array sensor for the simultaneous detection of MDA and 4-HNE
biomarkers.
In order to better understand and follow, at a nanostructured level, each step-to-step
modification along the assembly of the molecular imprinted film, other characterization
Chapter 8
129
methodologies can be employed in the future, such as, ellipsometry, AFM and additional
microscope strategies.
Additionally, the introduction of paper as the scaffold material of a biosensor device enabled
also the possibility to combine electrical and optical measurements simultaneously. Although
the electrochemical performance of the developed sensors was thoroughly explored, a different
option can be the incorporation of gold nanoparticles into the paper matrix, in order to develop a
facile colorimetric strategy. The main advantage of this kind of nanomaterials is that they can be
easily functionalized towards our target molecules and their aggregation behaviour can be
employed as colorimetric response.
In sum, the attributes of our biosensing approach can be compared to a very limited number of
other electrochemical devices, that are still using a conventional three electrode system, making
the paper-sustained device created herein the first electrochemical biosensor with potential to
become a portable and low-cost diagnostic tool towards OS biomarker detection.
Chapter 8
130
131
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